Fast thermo-optical particle characterisation

ABSTRACT

The present invention relates to a method and an apparatus for a fast thermo-optical characterization of particles. In particular, the present invention relates to a method and a device to measure the stability of (bio)molecules, the interaction of molecules, in particular biomolecules, with, e.g. further (bio)molecules, particularly modified (bio)molecules, particles, beads, and/or the determination of the length/size (e.g. hydrodynamic radius) of individual (bio)molecules, particles, beads and/or the determination of length/size (e.g. hydrodynamic radius).

The present invention relates to a method and an apparatus for a fast thermo-optical characterisation of particles. In particular, the present invention relates to a method and a device to measure the stability of molecules, like biomolecules, the interaction of molecules, particularly biomolecules with, e.g. further (bio)molecules, particularly modified biomolecules, particles, e.g. nanoparticles or microparticles, beads, e.g. microbeads and/or the determination of the length/size (e.g. hydrodynamic radius) of individual molecules, particularly of biomolecules, of particles (e.g. nanoparticles, microparticles), or of beads (e.g. microbeads) as well as the determination of e.g. length or size (e.g. hydrodynamic radius) of (bio)molecules or particles. Also combinations these characteristics may be determined with the means and methods of this invention. It is of note that the present invention is, however, not limited the measurement/characterization of biomolecules. Therefore, also the characteristics of other compounds/particles can be measured and determined by the means and methods disclosed herein, for example kinetic events and interactions of molecules may be determined and/or measured. Accordingly, also chemical reactions (like inorganic or organic reactions) may be measured by the methods and under use of the devices of the present invention. It is also envisaged to determine complex formations and/or their dissociation.

In conventional methods known in the art for all-optical biomolecule characterisation, samples with biomolecules in solutions are homogeneously heated to a certain temperature at a time followed by further heating to the next temperature point. A common procedure is to start at 20° C. The temperature is then increased for example by 1° C. A waiting time of approx. 2 minutes is then required until the whole system (cuvette and solution) has reached the applied temperature. This is due to the large thermal mass. Only then fluorescence is measured. This procedure is repeated until 90° C. in a stepwise manner. Accordingly, the heating of the whole sample volume takes long and it is necessary to employ heat conducting materials in contact with the liquid.

Separation techniques known in the art such as gel electrophoresis are at core of modern DNA and protein biotechnology. However, electrophoresis is hard to miniaturize due to electrochemical effects at the metal-buffer interface and the tedious preparation of the gel phase. Duhr et. al. in European Phys. J. E 15, 277, 2004 relates to “Thermophoresis of DNA determined by microfluidic fluorescence” and makes use of thermophoretic driving forces in miniaturized biotechnology devices. This article discusses an all-optical approach in thin micro fluids to measure and apply thermophoresis for biomolecules in small volumes. The temperatures are measured with high spatial resolution by the temperature sensitive fluorescence of a fluorescent dye. Typically, one measurement according to Duhr et al. (2004, loc. cit.) takes 300 s or even more. It is further speculated in Duhr et al. (2004, loc. cit.), that the movement of polymers, in particular DNA, in a temperature gradient is independent of the chain length of the molecule, an assumption in line with theoretical considerations, see e.g. Braun and Libchaber, Physical Review Letters 89, 18 (2002). This assumption strongly confines thermo-optical characterization of molecules based on thermophoresis, since the technique would solely be dependent on changes in size of molecules and would exhibit no sensitivity to surface properties, as it is the gist of the present invention.

The above mentioned method has the disadvantage that it is very time consuming. This is also the case for established methods for measuring interactions, size and stability, like Biacore (GE Healthcare), Evotec FCS-plus (Perkin-Elmer) or Lightcycler 480 (Roche Applied Science). The time consumption of these techniques is typically longer than an hour.

It is therefore an object of the present invention to provide an improved method and device for a thermo-optical characterisation of particles or molecules, in particular to provide a very fast method to measure thermally-induced processes of particles or molecules, in particular of biomolecules.

These objects are achieved by the features of the independent claims. Further preferred embodiments are characterized in the dependent claims.

The present invention relates in particular to a method and a device for measuring the stability of molecules, in particular biomolecules, the interaction of (bio)molecules with with other or further (bio)molecules, or with particles (e.g. nanoparticles, microparticles, beads, e.g. microbeads), and/or measuring the length or size of molecules, like biomolecules, particles (e.g. nanoparticles or microparticles, microspheres, beads, e.g. microbeads), Also combinations of these characteristics may be determined by the means and methods of this invention. The method of the present invention allows contact free, thermo-optical measurements of these parameters/characteristics within a time span of a few milliseconds up to a few seconds, i.e. a very fast analysis is possible. In the context of the present invention, a nanoparticle is a microscopic particle with at least one dimension less than 100 nm and a microparticle/microbead is a microscopic particle/bead which has a characteristic dimension of less than 1 mm but normally more than 100 nm.

In the context of this invention, in particular the claims, it is noted that the terms “particle” or “particles” also relate to beads, particularly microbeads, nanoparticles or molecules, particularly biomolecules, e.g. nucleic acids (such as DNA, RNA, LNA, PNA), proteins and other biopolymers as well as biological cells (e.g. bacterial or eukaryotic cells) or sub-cellular fragments, viral particles or viruses and cellular organelles and the like. The term “modified particle” or “modified bead” relates in particular to beads or particles which comprise or are linked to molecules, preferably biomolecules. This also comprises coating of such beads or particles with these (bio)molecules.

According to one embodiment of the invention, the inventive method is based on the absorbance of infrared LASER radiation by aqueous solutions and the subsequent conversion into heat. Thereby it is possible to create broad spatial, i.e., two-dimensional or three-dimensional (2D, 3D), temperature distributions comprising all desired temperatures, e.g. between 0° C. and 100° C. within an area of e.g. about 250 μm in diameter or length, established by local laser heating, whereby desired temperature gradients are created, in particular strong temperature gradients. Both, local temperature distributions and temperature gradients, can be used as described below to measure parameters, in particular biomolecular parameters. In a particular embodiment, the temperature distribution is on the micrometer scale. This may be advantageous since strong temperature gradients shorten the equilibration time of the system needs to equilibrate (i.e. measurement time). In particular embodiments, it is advantageous to increase the temperature on a length scale of less than 100 μm.

The present invention relates to a method to measure thermo-optically the characteristics of particles/molecules in a solution with the steps of: providing a sample probe with marked particles/molecules in a solution; exciting (e.g. fluorescently) said marked particles and firstly detecting and/or measuring (e.g. fluorescence of) said excited particles/molecules; irradiating a laser light beam into the solution to obtain a spatial temperature distribution in the solution around the irradiated laser light beam (i.e. in and/or nearby the area of the solution which is directly irradiated by the laser light beam); detecting and/or measuring secondly (e.g. a fluorescence of) the particles/molecules in the solution at a predetermined time after irradiation of the LASER into the solution has been started, and characterizing the particles/molecules based on said two detections.

Without differing from the gist of the invention it is also envisaged that instead of a detection based on fluorescence other detection methods are possible. Depending on the size and properties of the particles to be detected, the step of fluorescently exciting may be omitted, and a detection based on light scattering, (UV) absorption, phase contrast, phosphorescence and/or polarisation are possible. Moreover, for particles larger than 100 nm, the movement of such particles can be detected by single particle tracking.

The thermo-optical characterization in accordance with this invention allows to determine properties of molecules or particles in solutions, in particular in aqueous solutions. It also allows to discriminate between different conformations of one particle or molecule species and it also allows to discriminate between different species of particles or different molecules. The characterization can be used in all cases where the particles show a response to changes in the temperature gradient and changes in the absolute temperature. An advantageous feature of the present invention is the presence of a defined spatial temperature distribution. In particular, the temperature distribution is generated locally on microscopic length scales by local heating with a focussed laser. Another advantageous feature is that the response of the particles or molecules is assigned to a certain place of the known, optically generated spatial temperature distribution. Accordingly, temperature, place and response of the particles are directly correlated.

Furthermore, and in contrast to Duhr (2004; loc. cit.) the present invention provides means and methods for the thermo-optical measurements and/or thermo-optical characterization of particles or molecules, in particular biomolecules, by the measurements and/or the detection of differences in the thermo-optical properties. Their thermo-optical properties mainly originate from differences in thermophoretic mobility DT (i.e. the velocity of particles/molecules in a temperature gradient). In particular, the detected signal is dependent on the thermophoretic mobility c/c₀=exp[−(D_(T)/D)(T−T₀)], with the diffusion coefficient D, concentration c and temperature T. A DT independent of the polymer length as expected from Duhr (2004; loc. cit.) and others (e.g. Chan et al., Journal of Solution Chemistry 32, 3 (2003); Schimpf et al., Macromolecules 20, 1561-1563 (1987)) would render the analytics of biopolymers like DNA and proteins almost impossible since only changes in the diffusion constant would contribute to the thermo-optical properties, which are minute in most cases.

Thermo-optical characterization in accordance with this invention is based on the creation of strong temperature gradients at microscopic length scales in solution, in particular in aqueous solution. By doing so, the energetic states of the molecules in the solution are changed depending on the temperature and the properties of the molecule, i.e. the molecules experience a spatial potential originating from the spatial differences in temperature. This potential drives a directed motion of molecules along the temperature gradient, an effect called thermophoresis. In other cases the change in temperature leads, beside thermophoresis, to an unfolding of biopolymers, like proteins or DNA. The unfolding effect is observed at high temperatures and is a measure for the stability of molecules (the reason for unfolding is the increased influence of the entropy component of the energy), the effect is separated from thermophoresis by a characteristic time scale. The stability analysis takes place in milliseconds to one second, preferably about 1 ms to 250 ms, 1 ms to 200 ms, 1 ms to 100 ms, 1 ms to 80 ms, 1 ms to 50 ms, more preferably about 40 ms, 80 ms-180 ms, 80 ms-150 ms, most preferably about 50 ms.

Thermophoresis is observed at times in a range from about 0.5 seconds to 250 seconds, preferably 0.5 seconds to 50 seconds, preferably 1 second to 250 seconds, preferably 1 second to 50 seconds, preferably 1 second to 40 seconds, preferably 5 seconds to 20 seconds, preferably 5 seconds to 40 seconds, preferably 5 seconds to 50 seconds, preferably 5 seconds to about 80 seconds, more preferably 5 seconds to 100 seconds. Thermophoresis is a method which is sensible to surface properties of molecules in a solution. It is not necessary to expose molecules to a different matrix (like in chromatography) or to interact with the molecules physically in any way (e.g. by direct contact or by adding substances). Only interactions between electromagnetic waves and matter are necessary. Infrared radiation is used for spatial heating (i.e. manipulation of matter) and fluorescence to detect molecules.

The gist of thermo-optical characterization based on thermophoresis as provided herein is that differences in thermophoretic mobility (i.e. the velocity of molecules in a temperature gradient), and hydrodynamic radius can be detected by analyzing the spatial distribution of concentration (i.e. by the spatial distribution of e.g. fluorescence) or the fluctuations of single particles trapped in the spatial temperature profile. This embodiment is of particular relevance for the herein described thermo-optical trap for trapping particles, molecules, beads, cellular components, vesicles, liposomes, cells and the like. While the hydrodynamic radius is only related to the radius of a molecule, the thermophoretic mobility is sensitive to charge, surface properties (e.g. chemical groups on the surface), shape of a molecule (i.e. size of surface), conformation of a protein or interaction between biomolecules or biomolecules and particles/nanocrystals/microbeads. This means that if any of the mentioned properties are changed, the molecules will experience a different thermodynamical potential, resulting in differences in thermophoretic mobility (i.e. change in spatial concentration profile or fluctuation amplitude of trapped particles).

Thus, the present invention relates to thermally induced processes, e.g. temperature gradient induced directed motion or thermal denaturation.

The thermo-optical characterization mentioned above provides the means for fast thermo-optical analysis of particles and/or molecules, in particular for the thermo-optical characterisation of biomolecules, like nucleic acid molecules (e.g. DNA, RNA, PNA) or proteins and peptides. This characterisation comprises, inter alia, size determination, length determination, determination of biophysical characteristics, like melting points or melting curves, complex formations, protein-protein interactions, protein or peptide folding/unfolding, of intra-molecular interactions, intermolecular interactions, the determination of interactions between particles or molecules, and the like. Prior art methods for detection and quantification of molecular interactions and characteristics, in particular biomolecular interactions and characteristics are very time consuming which means that the time needed for an analysis is in the order of 30 minutes up to hours. In accordance with the present invention, one measurement typically takes less than 300 s, less than 200 s, less than 100 s or even less than 50 s, which is clearly faster than the methods described in the prior art. The present invention can detect and quantify molecular interactions and characterisations, in particular biomolecular interactions and/or biochemical/biophysical properties within 1 second to 50 seconds. The term interaction comprises interaction between biomolecules (e.g. protein, DNA, RNA, hyaluronic acids etc.) but also between (modified) (nano)particles/(micro)beads and biomolecules. In this context, modified particles, molecules, biomolecules, nanoparticles, microparticles, beads or microbeads comprise fluorescently labelled particles, molecules, biomolecules, nanoparticles, microparticles, beads or microbeads. Fluorescently labelled particles, molecules, biomolecules, nanoparticles, microparticles, beads or microbeads may be e.g. particles, molecules, biomolecules, nanoparticles, microparticles, beads or microbeads, to which one or more fluorescent dyes have been attached, e.g. covalently attached. For example, the fluorescent dyes may be selected from the group of 6-carboxy-2′,4,4′,5′,7,7′-hexachlorofluorescein (6-HEX SE; C20091, Invitrogen), 6-JOE SE or 6-TET SE (see also appended FIG. 6). In other cases, the intrinsic fluorescence of e.g. particles, molecules, biomolecules may be exploited according to the invention, e.g. the fluorescence properties of tryptophan, tyrosine or phenylalanine in a protein may be exploited. The terms “marked” and “labelled” are used synonymously in the context of this invention. In the context of this invention, “marked particles” refer to fluorescently labelled molecules/particles or other molecules/particles which can be detected by fluorescence means, e.g. molecules/particles comprising an intrinsic fluorophor, molecules/particles comprising intercalating dyes or particles/molecules with fluorophores attached.

A typical experiment in accordance with this invention, but not limiting the scope of the invention, to detect/quantify interactions may be described as follows:

Step 1a, Background Measurement:

A sample buffer without fluorescently labelled sample molecules/particles is filled into a microfluidic chamber and the fluorescence is measured, while the excitation light source is turned on.

Step 1b, Determination of Fluorescence Level Before Laser Heating:

An aqueous solution of a fluorescently labelled sample (e.g. biomolecules, particles, like nanoparticles or microparticles, beads, particularly microbeads, wherein in particular embodiments all of them have a specific affinity for other biomolecules) at a given concentration is filled in the microfluidic chamber (preferably a capillary) which preferably provides a defined height of the chamber. Fluorescence is excited and recorded with (e.g. CCD-Camera) or without (e.g. Photomultiplier tube, Avalanche Photodiode) spatial resolution for less than 10 seconds e.g. on a CCD device or photomultiplier with exposure times of 25 milliseconds up to 0.5 seconds. Then the fluorescence excitation is turned off.

Step 2, Starting of Infrared Laser Heating:

The infrared heating laser is turned on and the spatial temperature distribution is established within a few milliseconds within the solution. The temperature gradient has been calibrated once and it is not necessary to repeat this calibration every time an experiment is performed. In particular a setup where infrared heating and fluorescence imaging are performed through the same optical element from one side is advantageous for the stability of the optical and infrared foci.

A decrease of fluorescence due to photobleaching lower than 5% is advantageous in the experiment. In particular embodiments of the invention, no correction for photobleaching is necessary.

In some particular embodiments for measuring thermophoretic properties, the maximal temperature is below the temperature which is known to cause damage to the molecules in the solution or disturb their interaction (e.g. temperatures between 1 and 5° C. above ambient temperature).

Depending on the thermophoretic properties of the particles or molecules in the solution (i.e. if they move fast in a thermal gradient or slow) the infrared laser heats the solution for 5 seconds up to 100 seconds, preferably for 5 seconds up to 50 seconds, more preferably for 5 seconds up to 20 seconds.

Step 3, Recording of the Spatial Fluorescence (i.e. Concentration) Profile:

After this period of time, the fluorescence excitation is turned on and images are recorded with the same frame rate and length as described in step 1b. Step 3 is the last acquisition step necessary for evaluation of thermo-optical properties.

For detection and quantification of interactions more measurements following the protocol described previously are necessary. Step 1a is repeated with sample buffer and in step 1b the aqueous solution of a fluorescently labelled sample is mixed with an amount of the biomolecule with which the interaction should be detected or quantified. For example, in the detection of an interaction of particles and/or molecules, the fluorescently labelled sample (comprising one binding partner) is mixed with a sufficient amount of the second binding partner so that a substantial amount of the fluorescently labelled molecule or particle is in the complex with the binding partner. If the strength of the interaction is to be quantified in terms of e.g. a dissociation or association constant (Ka, Kd), than the procedure described previously may be conducted with varying concentrations of binding partner (e.g. 0.1×−10× the concentration of the fluorescently labelled binding partner). This means that a titration of binding partner may be performed.

Processing the Raw Data:

Optionally, a (linear) bleaching correction can be performed for which it is advantageous to wait for the back-diffusion of all molecules following the end of step 3. This increases the time consumption of the analysis dramatically. For precise and fast measurements it is advantageous to determine the strength of bleaching from image to image and correct every individual image with its own bleaching factor. For a precise bleaching correction it is advantageous that the temperature gradient at distance from the heat spot is low (e.g. below 0.001 K/μm). The images taken in step 1b are used to correct all images for inhomogeneous illumination. In case fluorescence is recorded without spatial resolution (e.g. avalanche photodiode or photomultiplier) photobleaching is corrected best by determining once the bleaching characteristic of a certain dye without heating laser in a control experiment.

Data Evaluation:

Qualitative detection of interaction: From the image series the spatial fluorescence distribution of the reference experiment (i.e. fluorescently labelled molecule/particle without binding partner) and the second experiment (i.e. were the binding partner is present) is extracted. The fluorescence is plotted versus the distance from the heat spot. An averaging is only possible for pixels with the same temperature and same distance. The spatial concentration distribution is obtained by correcting the fluorescence intensities for the respective temperature dependence of the dye. With knowledge of temperature dependence of the fluorescent dye and the spatial temperature distribution, the effect of a decreasing fluorescence due to temperature increase can be corrected. In particular embodiments, a correction for temperature dependency is not necessary for the qualitative detection of interaction as well as their quantification, and the spatial fluorescence distribution is sufficient. Any fluorescent dye on the market may be used, in particular embodiments even without characterization of its temperature dependency. The fluorescent properties of a dye may vary with buffer conditions, such as pH.

The values of the fluorescence profile are integrated up to the distance were the temperature is below e.g. 10% of the maximum temperature (e.g. 70 μm). The integrated values are compared and a change give a precise indication if there is an affinity between the substances at the concentrations used, since the interaction changes the thermo-optic properties (e.g. thermophoretic mobility, surface size and chemical groups on surface). In most cases, the interaction leads to higher fluorescence (concentration) at higher temperatures

In case the whole cross-section of a capillary is heated (i.e. using e.g. cylindrical lenses to give the IR laser beam a ellipsoidal shape, which heats a cross section of a capillary homogeneously), the intensity of two or more pixels from the centred heat spot may be averaged. In particular embodiments all pixels at the same distance to the heated line have the same temperature. This is advantageous for high precision measurements. In case fluorescence is recorded without spatial resolution, the fluorescence change in the centre of the heat spot/line is measured. In particular embodiments it may be advantageous to heat the whole cross section. In general if more than a single frame is recorded in step 1b and 3 an integration of multiple frames is possible.

For a quantification of molecular affinities or particle affinities the same procedure is performed for all experiments at various concentrations of non-fluorescent binding partner. The result of the integration for the reference experiment (i.e. without binding partner) is subtracted from the integrated values obtained for the different concentrations of binding partners. From this evaluation, the amount of interacting complexes in arbitrary units may be obtained. By dividing these values by the value where binding is saturated, the relative amount of formed complexes between the interacting molecules, particularly the binding partner, at a certain concentration of binder may be obtained. From these datasets also the concentration of free, e.g. non-fluorescent, binding partner may be determined and the strength of the interaction may be quantified in terms of association or dissociation constant (see also appended examples).

As mentioned previously, it is also possible to detect the binding of molecules to larger inorganic particles or nanocrystals using the procedure described previously and herein below. Inorganic particles, e.g. CdSe particles, may be modified with a varying numbers (e.g. 1, 2, 3, or 3 or more, yet preferably up to 3) of Poly-ethylen-glycol (PEG) of different molecular weight. In particular embodiments, 1 to 3 Poly-ethylen-glycol (PEG) molecules are attached to the particles. The spatial fluorescence profile is measured as described herein for the detection of biomolecular interactions (see appended examples). Also the raw data are processed as described herein. To measure the number or size of PEG molecules bound to the particles or nanocrystals, it may be sufficient to compare the spatial fluorescence profiles obtained with the protocol described previously. However, a correction for the temperature dependent decrease of the fluorescence allows a quantification in terms of the Soret coefficient. It is illustrated in the appended figures and examples, particularly in appended FIG. 26, that the Soret coefficient increases linearly with the number of PEG molecules bound to the nanocrystals. The slope of the increase depends on the molecular weight of the PEG. The binding of single molecules of the size of a protein is detectable as illustrated in e.g. appended FIG. 26.

The terms “interaction” or “affinity” as used herein and in particular in the above outlined, non-limiting illustrative experiment not only relates to the interaction of distinct molecules/particles (e.g. intermolecular interactions), but also to intramolecular interactions, like protein folding events and the like.

It is understood by the person skilled in the art that the term “fluorescence” as employed herein is not limited to “fluorescence” per se but that the herein disclosed means, methods and devices may also be used and employed by usage of other means, in particular, luminescence, like phosphorescence. Accordingly, the term “exciting fluorescently said marked particles and firstly detecting and/or measuring fluorescence of said excited particles” relates to the “excitation step” in the above identified method and may comprise the corresponding excitation of luminescence, i.e. excitation is carried out with a shorter wave length than detection of the following emission. Therefore, the term “detecting and/or measuring secondly a fluorescence of the particles” in context of this invention means a step of detection said emission after excitation. The person skilled in the art is aware in context of this invention that the “excitation”-wave length and the “emission”-wave length have to be separated.

According to a first illustrative embodiment of the present invention, the predetermined time (after which (e.g. a fluorescence of) the particles/molecules in the solution are detected and/or measured secondly) is small enough that concentration changes induced by thermophoresis and artefact related to or due to convection are negligible small. In other words, the predetermined time is short enough to separate the inter- and intra-molecular reaction time scale from slower temperature effects, e.g. thermophoresis, thermal convection. Thus, the predetermined time is preferably within the range of from 1 ms to 250 ms, more preferably between 80 ms and 180 ms, in particular 150 ms. In particular, in cases where the solution is provided in a chamber with good thermal conductivity, e.g. sapphire, diamond, and/or silicon, a shorter predetermined time is already sufficient, e.g. 1 ms, 5 ms, 10 ms or 15 ms. For the measurement it is advantageous that the chamber and the solution are in a thermal equilibrium. In other words, for chambers with a good thermal conductivity, the solution and the chamber are faster in a thermal equilibrium such that shorter predetermined times are sufficient. In case the thermal conductivity is poor, it takes longer until the chamber and the solution are in thermal equilibrium, i.e. the predetermined time is longer, e.g. 100 ms to 250 ms.

In particular embodiments of the invention, the detection or exposure time is in the range of from 1 ms to 50 ms. The time in which the detection signal is recorded must be short enough that the change of position of an individual molecule is negligible for the detection during the detection step. For example, in case the detection is conducted with a CCD camera with a resolution of 320×200 pixel, it is advantageous that during detection time an individual molecule/particle will be detected by only one pixel, since each pixel represents a certain temperature. If the position of a particle changes too much, i.e. more than one pixel, the particle may be exposed to a different temperature which decreases the measurement accuracy. The use of a CCD camera device for detection also comprises the use of a camera with only a single line of pixel (e.g. line camera) for one-dimensional detection.

In a particular embodiment of the invention, the laser beam is defocused such that a temperature gradient within the temperature distribution is in the range of from 0.0 to 2K/μm, preferably from 0.0 to 5K/μm. Thus, small temperature gradients ensure that the thermophoretical particle movement is negligible small during the time from the start of laser irradiation to the end of detection.

At least all temperatures needed to detect the thermal denaturation of a molecule have to be within the field of view of the camera device.

According to a further aspect of the present invention, the laser beam is irradiated through one or a plurality of optical elements into the solution. The focusing of the laser beam is in some embodiments performed in such a way that the temperature gradients lie within the above defined ranges. Focusing of the laser can be achieved e.g. by a single lens, a plurality of lenses or a combination of a optical fibre and a lens or a plurality of lenses or an objective where the divergence of the incident laser beam is adjusted properly. Moreover, further optical elements for controlling the focus and or direction of the laser beam may be arranged between the solution and the laser.

According to yet a further embodiment of the present invention, the temperature distribution around the laser beam is measured by an additional measurement, e.g. the temperature distribution is measured under the same conditions on the basis of the known temperature-dependent fluorescence of a dye, as illustrated in the appended figures, in particular FIGS. 3a, 3b and 15). In particular, a temperature distribution may be determined based on detected fluorescence of the temperature sensitive dye, wherein said temperature sensitive dye is heated (via the solution) by the irradiated laser beam and the fluorescence spatial fluorescence intensity is measured substantially perpendicular around the laser beam.

According to a second illustrative embodiment of the present invention, the predetermined time (after which (e.g. a fluorescence of) the particles/molecules in the solution are detected and/or measured secondly) is sufficiently long so that changes of concentration based on the thermophoretical motion can be detected. Thus, the predetermined time is preferably within the range of from 0.5 to 250 s. In said predetermined time the concentration changes within the spatial temperature distribution in the solution due to thermophoretic effects and such an concentration change may be detected by a change of the distribution of fluorescence.

In particular embodiments of the invention the laser beam is focused such that a temperature gradient within the temperature distribution is achieved in the range of from 0.001 to 10K/μm. The temperature within the field of view (particularly at the edge of the field of view) does not necessarily reach the value of ambient temperature. A temperature increase at distance to the heat centre (i.e. at the edge of the field of view) of 10% or less (in ° C.) in comparison to the maximal temperature (in ° C.) is advantageous.

According to a further embodiment, the fluorescence before and after irradiation the laser is detected with a CCD camera. The use of a CCD camera provides the advantage that the concentration change can be detected at a plurality of positions simultaneously. In particular embodiments, the CCD camera is a 2D (two dimensional) CCD camera, i.e., the CCD array comprises a plurality of sensor pixels (photoelectric light sensors) in a first and a second direction, wherein the first and second directions are preferably perpendicular to each other. According to a further embodiment, the CCD camera is a line or line-scan camera, i.e., the CCD array comprises a plurality of sensor pixels in a first direction (a line of sensor pixels) but merely one pixel in a second direction. Such a camera is also referred to as 1D (one dimensional) CCD camera. In other words, a one-dimensional array, used in line-scan cameras, captures a single slice or line of an image, while a two-dimensional array, captures a whole 2D image. The CCD array of the line camera may also comprise three sensor lines, each for one colour channel (Red, Green, and Blue). However, according to a further aspect of the present invention, it is also possible to measure the characteristics of the particles based on a detected fluorescence change of a single pixel of the CCD. Thus, it is also possible to use a photodiode or a photomultiplier instead of a single pixel of the CCD. In some embodiments the brightness of the fluorescence before and after irradiation the laser is measured with a photodiode or a single pixel with the CCD in the centre of the laser beam.

By imaging on a CCD camera device, a line camera or a PMT/Avalanche Photodiode, the fluorescence may be averaged throughout the height of the used liquid sheet. Accordingly, the three-dimensional solution may be reduced to two dimensions. Therefore, the method described in the embodiments is also applicable to two-dimensional lipid sheets, typically used as models system for membrane processes (e.g. surface supported tethered bilayer lipid membrane (tBLM, as illustrated in the appended figures, particularly FIG. 38), or classical Langmuir monolayers). Fluorescently labelled compounds floating in this membrane (e.g. lipids, proteins and alike) move in temperature gradients and rearrange according to their solvation energy. The fluorescence redistribution in these lipid layers or membranes may be employed in the context of the invention, like the redistribution of compounds in solution to detect biochemical or biophysical properties or characteristics, like conformational changes, interactions, hydrodynamic radius and the like. In the membrane system, the local temperature distribution is established by the surrounding aqueous solution, e.g. the aqueous solution above a surface supported membrane. By heat conduction the lipid layer above a aqueous solution also adopts the corresponding temperature.

According to the first and second illustrative embodiments of the invention, the particles can be biomolecules and/or (nano- or micro-)particles and/or beads, particularly microbeads, and combinations of those. With the use of modified nanoparticles/microparticles/microbeads proteins, DNA and/or RNA can be detected by specific bindings of proteins, DNA and/or RNA to the nanoparticles/microbeads, since the specific binding changes the thermophoretic motion of the nanoparticles/microbeads. The velocity of particles larger than 100 nm can be detected by single particle tracking.

The laser light may be within the range of from 1200 nm to 2000 nm. This range is advantageous for aqueous solutions. The hydroxyl group of water absorbs strongly in said wavelength range. Also other solvents with a hydroxyl group like glycerol can be heated by infrared laser heating). The laser is in some embodiments a high power laser within the range of from 0.1W to 10W, preferably from 1W to 10W, more preferably from 4W to 6W. In some embodiments the particle concentrations of an aqueous solution are within the range of from 1 atto Molar (e.g. single particle microbeads) to 1 Molar, preferably from 1 atto Molar to 100 μMolar.

According to a further embodiment, the solution may be a saline solution with concentrations in the range of from 0 to 1M.

According to still a further embodiment of the invention, the spatial temperature distribution generated by the LASER beam is in a range of 0.1° C. to 100° C. The temperature sensitivity of the material of interest sets the limits to the maximum temperature used in experiments. In particular temperature ranges of 0.1° C. to at least 40° C., preferably to at least 60° C., more preferably to at least 80° C. and even more preferably up to at least 100° C. are generated by the LASER beam to measure the DNA stability for example. The person skilled in the art is aware that corresponding temperatures can be achieved by, e.g., cooling of the experimental system as well as by use of LASER with corresponding power. The person skilled in the art is also aware that by cooling the overall sample a higher amplitude of temperature increase (i.e. by laser heating) is possible without causing damage to temperature sensitive materials. Illustrative, but not limiting, temperature ranges and maximum temperatures for different materials and thermo-optical characterizations is given in table 1. Accordingly, also higher temperatures can be achieved in the centre of the heat gradient (point of maximum laser power in and on the sample to be analyzed/characterized) as, inter alia, illustrated in appended FIG. 3. Such high temperatures may be achieved in high pressure chambers. Moreover, in some embodiments a temperature gradient is created within the range of from 0.1 μm to 500 μm in diameter around the LASER beam.

TABLE 1 Illustrative temperature ranges for thermo-optical analytics thermo-optic char.* conformation, Interaction, Hydrodynamic- structure, chemical Thermo-optic Sample Material Stability analysis kinetic events Radius, length modification Trap nucleic acids p. 0° C.-100° C. p. 0° C.-80° C. p. 0° C.-80° C. p. 0° C.-100° C. p. 0° C.-80° C. (DNA, RNA etc.) mr. p. 20° C.-95° C. mr. p. 20° C.-80° C. mr. p. 20° C.-80° C. mr. p. 10° C.-90° C. mr. p. 20° C.-60° C. mr. p. 30° C.-80° C. mr. p. 20° C.-40° C. mr. p. 20° C.-40° C. ms. p. 20° C.-90° C. mr. p. 20° C.-40° C. ms. p. 30° C.-85° C. ms. p. 20° C.-30° C. ms. p. 20° C.-30° C. ms. p. 20° C.-30° C. proteins p. 0° C.-100° C. p. 0° C.-40° C. p. 0° C.-60° C. p. 0° C.-80° C. p. 0° C.-40° C. mr. p. 20° C.-95° C. mr. p. 0° C.-30° C. mr. p. 0° C.-40° C. mr. p. 20° C.-60° C. mr. p. 10° C.-30° C. ms. p. 30° C.-85° C. mr. p. 10° C.-20° C. mr. p. 10° C.-40° C. mr. p. 20° C.-40° C. mr. p. 15° C.-30° C. mr. p. 15° C.-20° C. ms. p. 20° C.-30° C. ms. p. 0° C.-60° C. ms. p. 20° C.-30° C. ms. p. 20° C.-25° C. vesicles, liposomes p. 0° C.-100° C. p. 0° C.-40° C. p. 0° C.-60° C. p. 0° C.-80° C. p. 0° C.-40° C. mr. p. 20° C.-95° C. mr. p. 10° C.-40° C. mr. p. 0° C.-40° C. mr. p. 20° C.-60° C. mr. p. 10° C.-30° C. ms. p. 30° C.-85° C. mr. p. 20° C.-40° C. mr. p. 10° C.-40° C. mr. p. 20° C.-40° C. mr. p. 15° C.-30° C. ms. p. 30° C.-40° C. ms. p. 20° C.-30° C. ms. p. 20° C.-30° C. ms. p. 20° C.-30° C. ms. p. 20° C.-25° C. (Micro-)particles p. 0° C.-100° C. p. 0° C.-60° C. p. 0° C.-80° C. p. 0° C.-100° C. p. 0° C.-80° C. (e.g. silica, mr. p. 20° C.-95° C. mr. p. 10° C.-50° C. mr. p. 0° C.-60° C. mr. p. 10° C.-60° C. mr. p. 20° C.-60° C. polystyrene, etc) ms. p. 30° C.-85° C. mr. p. 20° C.-40° C. mr. p. 10° C.-50° C. mr. p. 20° C.-40° C. mr. p. 20° C.-40° C. ms. p. 20° C.-30° C. ms. p. 20° C.-30° C. ms. p. 20° C.-30° C. ms. p. 20° C.-30° C. PEG p. 0° C.-100° C. p. 0° C.-100° C. p. 0° C.-80° C. p. 0° C.-100° C. p. 0° C.-80° C. mr. p. 20° C.-95° C. mr. p. 10° C.-50° C. mr. p. 0° C.-60° C. mr. p. 10° C.-60° C. mr. p. 20° C.-60° C. ms. p. 30° C.-85° C. mr. p. 20° C.-40° C. mr. p. 10° C.-50° C. mr. p. 20° C.-40° C. mr. p. 20° C.-40° C. ms. p. 20° C.-30° C. ms. p. 20° C.-30° C. ms. p. 20° C.-30° C. ms. p. 20° C.-30° C. Sugar Polymers p. 0° C.-100° C. p. 0° C.-100° C. p. 0° C.-60° C. p. 0° C.-100° C. p. 0° C.-80° C. (e.g alginate, mr. p. 20° C.-95° C. mr. p. 10° C.-50° C. mr. p. 0° C.-50° C. mr. p. 10° C.-60° C. mr. p. 20° C.-60° C. hyaluroic acids) ms. p. 30° C.-85° C. mr. p. 20° C.-40° C. mr. p. 10° C.-40° C. mr. p. 20° C.-40° C. mr. p. 20° C.-40° C. ms. p. 20° C.-30° C. ms. p. 20° C.-40° C. ms. p. 20° C.-30° C. ms. p. 20° C.-30° C. two-dimensional p. 0° C.-100° C. p. 0° C.-40° C. p. 0° C.-60° C. p. 0° C.-80° C. p. 0° C.-40° C. lipidsheets (e.g. mr. p. 20° C.-95° C. mr. p. 10° C.-40° C. mr. p. 0° C.-40° C. mr. p. 20° C.-60° C. mr. p. 10° C.-30° C. containing ms. p. 30° C.-85° C. mr. p. 20° C.-40° C. mr. p. 10° C.-40° C. mr. p. 20° C.-40° C. mr. p. 15° C.-30° C. proteins) mr. p. 30° C.-40° C. ms. p. 20° C.-30° C. ms. p. 0° C.-60° C. ms. p. 20° C.-30° C. ms. p. 20° C.-25° C. (Nano-)particles p. 0° C.-100° C. p. 0° C.-100° C. p. 0° C.-80° C. p. 0° C.-100° C. p. 0° C.-80° C. mr. p. 20° C.-95° C. mr. p. 10° C.-50° C. mr. p. 0° C.-60° C. mr. p. 10° C.-60° C. mr. p. 20° C.-60° C. ms. p. 30° C.-85° C. mr. p. 20° C.-40° C. mr. p. 10° C.-50° C. mr. p. 20° C.-40° C. mr. p. 20° C.-40° C. ms. p. 20° C.-30° C. ms. p. 20° C.-30° C. ms. p. 20° C.-30° C. ms. p. 20° C.-30° C. Inorganic Carbon p. 0° C.-100° C. p. 0° C.-100° C. p. 0° C.-80° C. p. 0° C.-100° C. p. 0° C.-80° C. compounds (e.g. mr. p. 20° C.-95° C. mr. p. 10° C.-50° C. mr. p. 0° C.-60° C. mr. p. 10° C.-60° C. mr. p. 20° C.-60° C. carbon-nanotubes, ms. p. 30° C.-85° C. mr. p. 20° C.-40° C. mr. p. 10° C.-50° C. mr. p. 20° C.-40° C. mr. p. 20° C.-40° C. Buckyballs etc.) ms. p. 20° C.-30° C. ms. p. 20° C.-30° C. ms. p. 20° C.-30° C. ms. p. 20° C.-30° C. *The temperatures denoted here give the preferred temperature ranges of the sample at which thermo-optical properties may be measured. The temperature increase induced by laser heating may comprise only a small portion of the overall temperature, e.g. a thermo-optical characterization of nanoparticles can be carried out at 75° C. overall sample temperature with an increase in temperature by laser heating of 5° C.; a protein characterization may be carried out in a sample cooled to 10° C. and heated with an infrared laser to a maximum temperature of 30° C. p.: preferably, mr. p.: more preferably, ms. p.: most preferably In context of the invention, the term LASER is equivalent to the term “laser” and vice versa.

The present invention also relates to a device for measuring thermo-optically characteristics of particles in a solution. Such a device comprises a receiving means for receiving particles/molecules, particularly marked or labelled particles/molecules, within the solution, means for exciting the particles/molecules, particularly the marked or labelled particles/molecules, means for detecting the excitation of the particles/molecules, particularly the marked or labelled particles/molecules, and means for obtaining a spatial temperature distribution in the solution. A further device also referred to as device according to the present invention for measuring thermo-optically the characteristics of particles/molecules in a solution comprises a receiving means for receiving marked or labelled particles/molecules within a solution, means for fluorescently exciting the marked or labelled particles/molecules, means for detecting the excited fluorescence in said solution, and a laser for irradiating a laser light beam into the solution to obtain a spatial temperature distribution in the solution around the irradiated laser light beam. Particularly, the laser light may be focussed locally into the solution and the spatial temperature distribution may be obtained by the conduction of the absorbed energy as heat in the solution. By adjusting the focus width of the electromagnetic IR radiation, the spatial dimensions of the temperature distribution can be adjusted, i.e. a broad or a narrow temperature distribution is achieved. Further adjustments to the geometry of the temperature distribution can be achieved by choosing a material for the microfluidic chamber/capillary which exhibits a certain heat conductivity (e.g. high heat conductivity, narrow temperature distribution and vice versa). According to c/c₀=exp[(−D_(T)/D)(T−T₀)] the steady state amplitude of the thermo-optical signal is exponentially related to the increase in temperature. Using the relation shown above, a spatial temperature distribution can be fitted precisely to the spatial concentration distribution by adjusting the D_(T) and D coefficients. The time it takes the system to reach the steady state has, beside a dependence on D, also a strong dependency on D_(T) and the temperature gradient. The product of temperature gradient and the thermophoretic mobility D_(T) gives the velocity with which particles move along the temperature gradient. As a rule-of-thumb, the stronger the temperature gradient and the higher the thermophoretic mobility the shorter is the time needed to measure the thermo-optical properties. Therefore, it is advantageous that a temperature distribution is established on microscopic length scales (e.g. 250 μm) to obtain a strong temperature gradient.

According to further embodiments of the present invention, the means for exciting, preferably fluorescently exciting the particles/molecules or marked particles/molecules may be any suitable device selected from the group consisting of laser, fibre Laser, diode-laser, LED, Halogen, LED-Array, HBO (HBO lamps are, e.g., short arc lamps in which the discharge arc fires in an atmosphere of mercury vapour under high pressure), HXP (HXP lamps are, e.g., short arc lamps in which the discharge arc burns in an atmosphere of mercury vapour at very high pressure. E.g., in contrast to HBO lamps they are operated at a substantially higher pressure and they employ halogen cycle. HXP lamps generate UV and visible light, including significant portion of red light). Further means for exciting as referred to in the specification are also preferably used.

According to particular embodiments of the present invention, the means for detecting the excited particles, particularly for detecting the fluorescence, in the solution may be any suitable device selected from the group consisting of CCD camera (2D or line-scan CCD), Line-Camera, Photomultiplier Tube (PMT), Avalanche Photodiode (APD), CMOS-Camera. Further means for detecting as referred to in the specification are also used in particular embodiments.

The receiving means for receiving the particles, particularly the marked particles within a solution may be a chamber, a thin microfluidic chamber, a cuvette or a device for providing the sample in form of a single droplet. It is advantageous to provide the sample probe within a chamber which has a thickness in direction of the laser light beam from 1 μm to 500 μm, in particular 1 μm to 250 μm, in particular 1 μm to 100 μm, in particular 3 μm to 50 μm, in particular 5 μm to 30 μm. A person skilled in the art will understand that the term chamber also relates to e.g. a capillary, microfluidic chip or multi-well plate. In some embodiments the chamber has the same width as the dimension in direction of the laser light (e.g. a capillary). In combination with a corresponding ellipsoidal laser heating geometry such a system can be reduced from radial symmetry to a single dimension since a CCD camera with only a single line of pixels (line-scan CCD) can be used to integrate the fluorescence of the whole width of the chamber. In a particular embodiment, receiving means for receiving the marked particles within a solution alias specimen holder is attached to an optical element, such as an objective. Such a setup avoids relative movements of the specimen holder with respect to the objective. Further means for receiving as referred to in the specification are also used in some embodiments.

The laser for irradiation the laser light may e.g. be an IR laser, e.g., a laser with a wavelength between 1200 and 2000 nm, preferably of 1455 nm and/or 1480 nm and a radiation power of 0.1 to 10W. The light of the Laser may be coupled into the device of the present invention by means of optical units like laser fibres (single mode or multimode) with or without collimator. Further means for irradiating as referred to in the specification are also used in some embodiments and are within the normal skill of the artisan.

The device according to the present invention may further comprise a control unit for controlling the means for exciting the particles and/or the means for detecting the excited particles. In particular, the control unit is adapted to enable the device of the present invention to perform the method steps as discussed with regard to the present inventive method.

The control unit may control the type (e.g. wavelength), the intensity, the duration, and/or the time for start and stop of the irradiation of the means for the excitation. For instance, in a particular embodiment in which the means for excitation is a laser, the duration and/or the start and stop time for generating a laser beam may be controlled by the control unit.

The control unit may further or alternatively control the exposition, the sensitivity, the duration, and/or the time for start and stop time for detecting/measuring by means of the means for detecting. For instance, in a particular embodiment in which the means for detecting is a CCD camera, the exposition timing of the CCD camera may be controlled by the control unit.

The control unit may be further adapted to control the means for detecting dependent on functional state of the means for exciting. In particular, it may be advantageous to synchronize the timing of excitation with the timing of the detection. For instance, in a particular embodiment in which the mean for excitation is a laser and the means for detecting is a CCD camera, the exposure timing of the CCD is synchronized with the irradiation of the laser. This may be achieved by controlling the CCD and the laser directly, e.g. by switching on and off the CCD and the laser synchronously.

The control unit may alternatively or additionally control means, in particular optical means, which are arranged between the means for the excitation and the receiving means and/or are arranged between the means for the detecting and the receiving means.

The device according to the present may comprise at least a shutter arranged between the receiving means for receiving the marked particles/molecules and the means for exciting the particles/molecules, in particular the laser. The device according to the present may additionally or alternatively comprise at least a shutter arranged between the receiving means for receiving the marked particles and the means for detecting the particles/molecules, in particular the CCD camera. Such a shutter may be controlled by the control unit in order to adapt the timing of the step of excitation with the timing of the detection.

A single control unit may fulfil the functions of several items, i.e. the control unit may comprise a plurality of subunits which are adapted for controlling particular means.

The device according to the present may further comprise at least a beam splitter and/or a mirror, e.g. a dichroic filter or dichroic mirror, i.e. a colour filter used to selectively pass light of a small range of colours while reflecting other colours, or a AOFT. The dichroic mirror may reflect short wavelength (reflectance>80%) and transmit long wavelength (transmission >80%). The dichroic mirror may e.g. comprise an IR-transmission larger than 90% and a at least a reflectance for wavelength between 350 to 650 nm. In some embodiments, instead of a dichroic mirror a silver mirror may be used. The (dichroic) mirror may be arranged in a fixed position within the device. However, according to some embodiments, the (dichroic) mirror may be movable, e.g. driven by a driving means (which may be controlled by the control means).

The device according to the present may comprise at least an emission and/or excitation filter (band pass/long pass) for filtering specific wavelength.

In other words, the device according to the present invention may also comprise optical means arranged between the receiving means for receiving the marked particles and the means for detecting the particles and/or between the receiving means for receiving the marked particles and the means for exciting the particles. Such optical means may be adapted for controlling the propagation direction of light by way of transmission or reflection and/or for filtering or separating different wave length by (dichroic) filters. Such optical means may be passive optical means or active optical means which may be controlled by the control unit. For instance, there may be arranged a scanning module (e.g. a Galvano scanning mirror) between the receiving means for receiving the marked particles and the means for exciting the particles. The scanning range and timing of such a scanning module may be controlled by the control unit, preferably in dependency with the control of the means for receiving and/or means for exciting.

Optical means which are advantageous for the device according to the present invention, in particular to perform the method steps of the present invention are exemplified above and below. In particular a plurality of optical means which are useful for a device of the present invention are illustrated in the detailed description of the invention.

According to a further embodiment of the invention, the irradiation of the laser and the detection of the fluorescence is conducted from different directions, e.g. the irradiation is from below and the detection is from above the sample (as illustrated in the appended figures, particularly FIG. 1). However, the means for irradiation and the detection can be arranged on the same side with respect to the sample probe (see e.g. FIG. 2). The device according to the present invention may have any orientation with respect to the direction of gravity, i.e., the device may e.g. be oriented substantially perpendicular, parallel or anti-parallel with respect to the direction of gravity.

In particular embodiments of the invention the sample probe is provided in a chamber. The thickness of the sample probe in the chamber in direction of the laser light beam is preferably small, e.g. from 1 μm to 500 μm, in particular 1 μm to 250 μm, in particular 1 μm to 100 μm, in particular 3 μm to 50 μm, in particular 5 μm to 30 μm. A person skilled in the art will understand that the term chamber also relates to e.g. a capillary, microfluidic chip or multi-well plate. In a further embodiment the chamber has the same width as the dimension in direction of the laser light (e.g. a capillary). In combination with the corresponding ellipsoidal laser heating geometry such a system can be reduced from radial symmetry to a single dimension. A CCD camera with only a single line of pixels can be used to integrate the fluorescence of the whole width of the chamber. Furthermore only a single pixel (or a photodiode or photomultiplier) mapped to the centre of the heat spot can be used for detection of interactions, conformation etc. It is illustrated in the appended figures, particularly FIG. 27, how a capillary is used for thermo-optical characterization. A capillary may be placed on a solid support/specimen holder/stage with good heat conducting properties. The solid support/specimen holder/stage may be cooled or heated by a Peltier element. By using a Peltier element the “ambient temperature” of the solution can be adjusted. It is advantageous to measure the protein conformation at different temperatures, to adjust thermophoresis of (bio)molecules/(nano)particles/(micro)beads to a value close to the sign change of thermophoresis (i.e. that a binding event changes the thermophoretic behaviour from accumulation at higher temperatures to depletion from higher temperatures). A cooling of the chamber allows to heat temperature sensitive molecules with a much higher laser power. In further embodiments valves are put at the end of the capillary to exclude any drift of the liquid inside the capillary. Often drift is caused by evaporation at the end of a capillary.

However, it is also possible to provide the sample probe without a chamber such as in form of a droplet, e.g. buffer droplet.

In some embodiments of the invention the fluorescence is detected within a range of from about 50 nm to 500 μm in direction of the laser beam.

In further embodiments the fluorescence is detected substantially perpendicular with respect to the laser light beam with a CCD camera. The second fluorescence detection is in some embodiments a spatial measurement of the fluorescence in dependence of the temperature distribution substantially perpendicular with respect to the laser light beam.

The appended figures show particular, non-limiting setups for devices in accordance with the present invention in accordance with the present invention. These devices are particularly useful in measuring thermophoresis. Common to all of them is that fluorescence imaging and the infrared laser focussing are performed through the same optical unit, e.g. through the same objective. Particularly, an objective with very low refraction of the IR light may be used. This is beneficial because a local spatial temperature profile on the micrometer scale is advantageous and desired in context of the means and methods provided herein. High refraction of IR light by the objective may lead to a broad temperature distribution with a high background temperature elevation. To solve this less advantageous effect, an objective with high transmission in the IR region of the electromagnetic spectrum (preferably at 1200-1600 nm, i.e. corrected for IR radiation) may be used. An objective which comprises only a small number of lenses (i.e. objective with less correction for visible wavelengths) is here preferred. As described herein, if a microfluidic chamber with a high aspect ratio (length/width) is used it may be advantageous to change the IR laser beam profile to an ellipsoidal form, to homogeneously heat the whole cross section of the capillary. This allows, inter alia, high precision measurements with a line camera, photodiode or photomultiplier. A line camera provides merely line resolution along the capillary and averages the spatial fluorescence of a whole cross section (width) of the capillary.

The photodiode or photomultiplier has no spatial resolution, but it is positioned in a way to measure the fluorescence in the central heated region. Such a positioning is within the normal skills of the person skilled in the art. Both, line camera and photodiode combined with an ellipsoidal illumination of the microfluidic chamber (i.e. capillary) are used in some embodiments of the invention for data acquisition.

Appended FIG. 23 illustrates a corresponding further embodiment, wherein the simultaneous detection of two or more different fluorophores/marked particles is documented. The different emission wavelength of the two or more marked particles/molecules are splitted e.g. via a dichroic mirror or AOTF into two or more different directions. In this embodiment, one detection channel can for example be used to measure the temperature via the temperature dependent fluorescence of a dye, e.g. Cy5 at a wavelength for example 680 nm+/−30 nm. In the other channel the melting curve of a marked particle/molecule can be recorded at a wavelength of for example 560 nm+/−30 nm. It also allows for the parallel detection of particles/molecules, e.g. different particles/molecules, with e.g. different luminescent or fluorescent markers.

An advantage of the present invention, is that, in contrast to the prior art, now particles, in particular and as exemplified, (bio)molecules or (nano- or micro-)particles or (micro)beads can be measured/determined/characterized by employing spatial temperature distributions with μm resolution.

Accordingly, with the means, methods and devices provided herein it is, inter alia, possible to measure, detect and/or verify biological, chemical or biophysical processes and/or to investigate, study and/or verify samples, like biological or pharmaceutical samples. Also diagnostic tests are feasible and are embodiments of this invention. It is, inter alia, envisaged and feasible to measure the length of nucleic acid molecules (like DNA, RNA), to measure the melting features of proteins or nucleic acid molecules, like, e.g. of double-stranded DNA or double-stranded RNA (dsDNA/dsRNA) or of hybrid nucleic acid molecules, like DNA/RNA hybrids, to measure and or analyze nucleic acid sequences, like the detection and/or measurement of Single Nucleotide Polymorphisms (SNPs) (see also appended figures, in particular FIG. 4) or to measure the stability of nucleic acid molecules in correspondence and as a function of their relative length; to measure and/or verify PCR end products, e.g. in general medical diagnostic, also in polar-body diagnostic, pre-implantation diagnostic, forensic analysis. Accordingly, it is evident for the skilled artisan that the means and methods provided in this invention are, particularly and non-limiting, useful in measurements and/or verifications wherein the length, size, affinities to other molecules/particles of a given particle/molecule is of interest. For example, the methods provided herein as well as the devices are useful in the detection and measurement of length, temperature stability as well as melting points of nucleic acid molecules and proteins. Therefore, it is within the scope of the present invention that, for example, (DNA-) primers and (DNA- or RNA-) probes are measured and or verified after or during their synthesis. Also the measurement of nucleic acid molecules on templates, like DNA-chips is envisaged. The term melting in context to this invention refers to the thermal denaturation of biomolecules, like nucleic acids (e.g. RNAs, DNAs) or proteins.

Also envisaged in context of the present invention is the measurement, detection and/or verification of mutations and genetic variations in nucleic acid molecules, for example in form of single-strand conformational polymorphisms (SSCPs) or in form of restriction fragment length polymorphisms (RFLPs) and the like. The present invention also provides for the possibility to analyze heteroduplexes. Heteroduplexes are generated by heat denaturation and reannealing of a mixture of e.g. wild type and mutant DNA molecules. In particular it is also possible to measure the effect of protein binding to a DNA molecule on the stability of the latter. Furthermore it is possible to measure the thermal stability of proteins and the effect of molecules (e.g. small molecules, drugs, drug candidates) on the thermal denaturation.

Also within the scope of the present invention is, e.g., the measurement of protein-protein interactions, like complex formations of proteinaceous structures or of proteins or of fragments thereof. Such measurements comprise, but are not limited to the measurement of antibody-antigen binding reactions (also in form of single chain antibodies, antibody fragments, chromobodies and the like). Yet, the embodiments of the present invention are also related to the detection and or measurement of dissociation events, like, e.g. the dissociation of protein complexes. Therefore, the invention is also useful in the measurement, determination and/or verification of dissociation events, like in the measurement of the dissociation of proteinaceous complexes, e.g. antibody-antigen complexes and the like. The appended figures also show that the means and methods provided herein are useful, for example, in the measurement of melting curves for nucleic acid molecules, proteins and corresponding analyses.

It is understood by the person skilled in the art that the term “modified microparticle/nanoparticle” as employed herein is not limited to “microparticle/nanoparticle” per se but that the herein disclosed means, particles and materials may also be used and employed by usage of other means, in particular, colloidal means, like foams, emulsions and sols.

As microparticles show a very strong thermophoresis, as illustrated in the appended examples and figures, particularly in FIG. 33, they can be used as a carrier material for the detection and characterisation of for example biomolecules, like proteins or nucleic acids. By using microparticles, the thermophoresis signal of e.g. the biomolecules may be enhanced. The attraction of a microparticle to an extreme value (e.g. the temperature maximum) of the spatial temperature distribution due to thermophoresis may be strong enough to trap the microparticles there (as illustrated in the appended examples and figures, particularly in FIG. 34).

A microparticle is a particle with a characteristic length of less than 1 mm and more than 100 nm without restriction of the material (e.g. coated or uncoated silica-/glass-/biodegradable particles, polystyrene-/coated-/flow cytometry-/PMMA-/melamine-/NIST particles, agarose particles, magnetic particles, coated or uncoated gold Particles or silver Particles or other metals, transition metals, biological materials, semiconductors, organic and inorganic particles, fluorescent polystyrene microspheres, non-fluorescent polystyrene microspheres, composite materials, liposomes, cells and the like).

A nanoparticle is a particle with a characteristic length of less than 100 nm without restriction of the material (e.g. quantum dots, nanocrystals, nanowires, quantum wells).

Particles or beads according to this invention may be modified in such a way that for example biomolecules, e.g. DNA, RNA or proteins, may be able to bind (in some embodiments specifically and/or covalently) to the particles or beads. Therefore, within the scope of this invention is the thermo-optical analysis of characteristics of beads and/or particles and in particular of molecules attached to or linked to such beads or particles. In particular, such molecules are biomolecules. Accordingly, the term “modified (micro)beads/(nano- or micro)particles”, in particular, relates to beads or particles which comprise additional molecules to be analyzed or characterized (a non limiting example for the case of nanoparticles is shown in the appended figures and examples, particularly in FIG. 34). Modified or non-modified microparticles/(nano- or micro)particles may be able to interact with other particles/molecules such as biomolecules (e.g. DNA, RNA or proteins) in solution. The skilled person understands that the thermophoretic properties of the modified particles will change upon binding of the biomolecules in solution to the biomolecules bound to the particle as modification. Such an interaction may influence the force acting on the (modified) particle/molecule. By adjusting the IR-Laser irradiation, the resulting movement can be influenced in such a way that the particle/bead is trapped. “Trapping” the particle/bead, in particular particles/beads comprising biomolecules, means that the particle/bead stays within a certain position, only showing comparably low fluctuations. These fluctuations are different from fluctuations based on Brownian motions. When a biomolecule from solution binds to the biomolecular-modified particle, the force acting on the particle/bead will change due to a change of the thermophoretic properties, which may result in movement of the particle out of the certain position where the particle/bead was trapped or/and in a change of the fluctuations of the particle/bead. The method described here is termed “Thermooptical Trap” and is particularly useful in certain embodiments described herein. The “Thermooptical Trap” is also illustrated in the appended examples and figures. Other synonyms for “Thermooptical Trap” are “Optothermal Trap”, “Thermophoretic Trap” as well as “Optothermal Tweezers” or “Thermophoretic Tweezers”.

The invention, accordingly, also relates to an optothermal trap. The terms “thermo-optical”, “thermooptical”, “optothermal” and “opto-thermal” are used synonymously. This particular embodiment of the invention illustrates that given targets, e.g. fluorescently labelled modified beads/particles with a size of 100 nm up to several μm (e.g. polystyrene beads or silica beads) and lipid vesicles and cells are showing a directed movement to the temperature maximum of a spatial temperature distribution generated by applying IR-Laser radiation to an aqueous solution as illustrated in the appended figures, particularly FIG. 39.

As shown in the appended example, the devices and the methods of the present invention may also be employed for thermophoretic trapping of molecules or particles, including lipid structures (like vesicles or liposomes) as well as cellular components of even cells. It is also envisaged that the devices and methods of the present invention is used for thermophoretic trapping of cells or cellular components, like cell nuclei, chromosomes, mitochondria, chloroplasts and the like. The thermophoretic trapping as shown herein is particular useful for studying interactions of e.g. proteins (e.g. with other proteins, for example antibody-antigen interactions and the like), for studying transport events across membranes (e.g. vesicles or liposomes), for determination of activity of membrane proteins comprised in biological membranes/vesicles/liposomes, like ion pumps, membrane transporters and the like. Also, the mere presence of molecules, particles, liposomes, vesicles, beads, cells or cellular components in said solution can be detected and/or analyzed by use of the herein disclosed thermophoretic trapping devices and methods. Thermophoretically trapped molecules, particles, vesicles, beads, cells or cellular components and the like can be transported and moved within the analyzing solution (see also appended figures). It is envisaged that thermophoretically trapped molecules, particles, liposomes, vesicles, beads, cells or cellular components be exposed to different buffer solutions for certain applications, i.e. buffer around the trapped molecules, particles, vesicles, beads, cells or cellular components etc. may be exchanged and corresponding measurements can be carried out. Further embodiments in context of the thermophoretic trapping in accordance with this invention are also provided in the appended examples. It is evident for the person skilled in the art that the concepts of thermophoresis as disclosed herein can also be employed for example in sorting of vesicles, cellular components (like e.g. mitochondria, chloroplasts, nuclei, chromosomes) or even whole cells. Accordingly, the present invention also provides for a method to thermo-optically trap molecules, particles, vesicles, beads, liposomes, cells or cellular components etc., said method comprising the steps of providing a sample probe with (preferably marked) molecules, particles, vesicles, beads, liposomes, cells or cellular components; irradiating a laser light beam into the solution to obtain a spatial temperature distribution in the irradiated laser light beam; optionally detecting the (preferably marked) molecules, particles, vesicles, beads, cells or cellular components; and trapping the (preferably marked) molecules, particles, vesicles, liposomes, beads, cells or cellular components in accordance to the thermophoretic mobility of said molecules, particles, vesicles, beads, cells or cellular components. For example, the (preferably marked) molecules, particles, vesicles, beads, cells or cellular components can be trapped in center of the laser-generated heat spot (in particular when thermophoretic mobility of the trapped molecules, particles, vesicles, liposomes, beads, cells or cellular components etc is negative). However, the (preferably marked) molecules, particles, vesicles, beads, cells or cellular components can also be trapped in a (global or local) temperature minimum, in particular, when the thermophoretic mobility of the trapped molecules, particles, vesicles, beads, cells or cellular components etc is positive. The person skilled in the art is aware that the term “thermophoretic mobility”, DT refers to a coefficient which relates the velocity (v) of a given molecule/particle/bead etc to the temperature gradient (∇T), according to v=−DT∇T.

The embodiments provided herein above in context of the method for measurement of thermo-optical characteristics of particles/molecules etc. in a solution apply to the method of thermo-optically trapping molecules, particles, vesicles, beads, liposomes, cells or cellular components, etc mutatis mutandis. Also devices for the thermo-optically trapping are provided herein and are also illustrated in the appended figures, for example the device as shown in appended FIG. 19 or 24. A corresponding device comprises, accordingly, an IR laser for irradiating a laser beam into the solution containing the (preferably marked) molecules, particles, vesicles, beads, liposomes, cells or cellular components to be trapped, to obtain a spatial temperature distribution in said solution around the irradiated laser light beam. Said device for thermo-optically trapping (preferably marked) molecules, particles, vesicles, beads, liposomes, cells, cellular components, etc, may therefore comprise: (a) a receiving means for receiving (optionally marked) molecules, particles, vesicles, beads, cells, cellular components, etc within a solution; (b) (optionally) means for fluorescently exciting the marked particles; (c) (optionally) means for detecting the excited fluorescence in said solution; and (d) an IR laser for irradiating a laser light beam into the solution to obtain a spatial temperature distribution in the solution around the irradiated laser light beam.

The movement of molecules, particles, vesicles, liposomes, beads, cells, cellular components, etc. can be described by a thermophoretic force acting on the particle. Presuming thermodynamic equilibrium, this force may be derived from the Gibbs free enthalpy at constant pressure: F=−½*S _(T) *k _(B)*grad(T ²) Where S_(T) is the Soret coefficient and k_(B) the Boltzmann constant.

The temperature T is a function of x and y: T=T(x,y). For example, in a radial symmetric geometry like the spatial temperature distribution generated by a focused IR-laser (as illustrated in the appended figures, particularly in FIG. 3a ) it can easily be seen that the force can attract a particle to the maximum of the spatial temperature distribution. If S_(T)<0 this results in a force trapping the given target particle/bead, e.g. a silica microparticle, preferably a coated silica microparticle, at the maximum (local or global) of the temperature distribution.

In contrast to the optical tweezers/optical trapping known from the art, the thermo-optical trap according to the present invention is based on different principles. Instead of using an electromagnetic field gradient as used for optical tweezing, a temperature gradient is used to trap, move and control a particle according to the present invention. Therefore, no sophisticated confocal optic is necessary in some embodiments of the thermo-optical trap. Using temperature gradients also allows for attraction of molecules from distances of 1 μm to several hundred micrometer distance to the laser focus, depending on the width of the temperature gradient (e.g. the width of the IR laser focus). Compared to the thermo-optical trap, the catchment region for optical tweezers is very narrow on the order of a few micrometer.

As shown in the appended Example 1, the Soret coefficient S_(T) is a function of the surface area A of the target particle, e.g. the bead, the quadratic effective charge σ_(eff) and particle-area-specific hydration entropy s_(hyd). So the thermophoretic force will also be proportional to this particle properties.

If one of this properties changes (preferably the effective charge or the hydration entropy), the trapping force changes also. If the trapping force/the trapping potential changes, the fluctuations (as also illustrated in the appended FIG. 32) also change By recording the fluctuations of the particle, a change in the thermophoretic properties of the particle can be detected and so the binding of for example biomolecules to this particle/bead can be detected.

In case of a particle/bead, e.g. a silica microparticle, preferably a coated silica microparticle, more preferably a silica microparticle coated with special groups on its surface (e.g. proteins). These groups may be able to specifically bind to proteins, antibodies, small molecules, DNA, RNA etc. If there is a binding of one of this species to the specific group on the bead/particle, the properties (for example the surface A) of the particle, e.g. bead changes, which may result in a different S_(T) and thus to a different thermophoretic force F (e.g. the sign/direction of the force may change).

A particular embodiment of the invention relates to the measurement of the fluctuations of such a given target particle/molecule in the maximum of the temperature distribution by detecting its position via measurement of fluorescence or the like. When the particle/molecule is trapped in the temperature distribution, the solution around it can be easily exchanged with a solution containing non-labelled or labelled molecules which are envisaged to bind to the “trapped” particle, e.g. when modified in a way that allows the specific binding (e.g. by antibodies). A binding event may be detected by a change in the amplitude of fluctuations (i.e. the potential in which the particle is trapped changes due to the binding of molecules to the particle's surface). By detecting the time dependent characteristics of the variation in amplitude, a binding kinetic can be measured and established.

In a further particular embodiment of the invention, the change in sign of S_(T) of the target particle may be exploited. If S_(T)<0 a particle, e.g. a bead or a fluorescent silica microparticle, with specific binding groups/sites is trapped at the maximum temperature in the spatial temperature distribution. If S_(T) of the target particle changes its sign upon e.g. crowding or linking of the binding groups with molecules e.g. small molecules, the target particle is forced away from the maximum of the spatial temperature distribution and the particle, e.g. bead, experiences repulsion instead of attraction, resulting in a detectable qualitative change of the behaviour of the target particle. For example if the force changes from attraction to repulsion, the particle may move away from the maximum of the spatial temperature distribution. Thereby, molecules, e.g. small molecules, binding to said binding groups at the surface of the particle/bead, may be detected. The binding of molecules, e.g. small molecules, can easily be measured just by simple particle tracking methods. The particle can be brought to conditions close to a sign change by changing the buffer conditions (e.g. salt concentration) the temperature of the solution or by specific modifications with e.g. hydrophobic or charged molecules. It is evident that the term “target particle” or “target bead” also relates in this embodiment to correspondingly modified particles and beads, like particles/beads comprising or linked to biomolecules or particles/beads coated with such biomolecules. Therefore, the above recited embodiments for “particles/beads” and in particular modified particles/beads apply here mutatis mutandis.

The force acting on the target particle, e.g. a silica microparticle coated with special groups, can be measured by tracking its position by an appropriate method like fluorescence (if the particle is fluorescently excitable), phase contrast, interference, farfield imaging etc.

It is also possible to apply a second force onto the target particle, e.g. a specifically coated magnetic microparticle, in a way that the resulting force is a superposition of the thermophoretic force and the second force (e.g. a magnetic force). The second force can be for example a magnetic force (for magnetic particles) or an electric force (for charged particles) or a hydrodynamic force or another optical force like it is generated by an optical tweezers.

Then the resulting superposition of forces may then be measured, for example by measuring the fluctuations of the particle (as also illustrated in the appended FIG. 32). This superposition may be used to increase the sensitivity of the method for example by using counteracting forces in such a way that a small change in one of the properties of the target particle/bead, will result in a movement of the target particle/bead, whereas without a change in the properties of the particle/bead, the particle/bead stays at the same place. The increase of sensitivity may be due to the possible minute adjustability of the second force (e.g. a magnetic force).

The “Thermooptical Trap” can also be used to move a target particle/molecule, e.g. a bead/particle, in two dimension perpendicular to the axis of the incident IR-Laser radiation. If the focus of the IR-Laser is moved for example by use of galvanic mirrors or an acoustic optic deflector (AOD) the resulting maximum of the spatial temperature distribution is also moved and so is the target particle/molecule/bead. Or vice versa, the chamber may be moved and the IR-Laser focus is kept fixed (as also illustrated in the appended figures, particularly FIG. 34).

Using a variety of IR-Laser focal spots lots of particles/beads/molecules can be moved simultaneously, giving the opportunity of multiplexing and also of combining different target particles, e.g. specifically coated microparticles, with each other. So if there is one target particle with an antibody and another target particle with the corresponding antigen, the target particles can be moved by two IR-Laser foci until they are in contact and the antibody binds to the antigen. In this way the target particles/molecules bind to each other and a composition of particles may be created.

In accordance with this invention, it is also possible to generate an interference pattern of the IR-Laser radiation, resulting in a spatial grating of temperature maxima. With this spatial grating of spatial temperature distribution target particles/molecules can be trapped and they can also be moved by moving the interference pattern.

As also demonstrated in the appended figures, the present invention is particularly useful in the determination of single or double stranded nucleic acid molecules (see, e.g. appended FIG. 5). This allows, inter alia, to determine in a given probe/sample whether it comprises single and/or double stranded nucleic acid molecules. This is in particular relevant in cases when it has to be determined whether a given biological sample comprises, e.g. viral nucleic acids, like single stranded DNA or single stranded RNA.

As documented herein, one embodiment of the present invention is based on the fact that with the means and methods of this invention it is possible to measure, in very short time intervals, inter- as well as intra-molecular interactions. In the first illustrative embodiment of the invention as described herein, a thermo-optical method is disclosed that allows for the detection of a broad temperature range (in a given probe/sample) at the same time, whereby said “same time” is a time range of about 1 ms to 250 ms, in particular 80 ms to 180 ms and, as exemplified at 150 ms, yet, at the most at 250 ms. This first illustrative embodiment of the invention is not based on or related to thermophoresis. In contrast, thermophoresis is, to a large extend excluded. The first embodiment is, e.g. related to the determination of melting curves, e.g. the determination of DNA and protein melting (point) curve(s). A non-limiting example of this first embodiment is the determination/measurement of single nucleotide polymorphism, based as provided in appended FIG. 4. The “melting point” is defined by 50% dissociated molecules. It is evident herein that the disclosed method as provided in the first embodiment is not limited to the determination of melting points of DNA molecules.

In the second illustrative embodiment of the invention, thermophoresis or thermophoretic effects play a role, inter alia, in a pre-determined time of about 0.5 second to about 250 seconds, preferably about 1 second to about 150 seconds, more preferably about 5 seconds to about 100 seconds, more preferably about 5 seconds to about 80 seconds more preferably about 5 seconds to about 50 seconds and even more preferably in about 5 seconds to about 40 seconds, concentration changes within a spatial temperature distribution are measured and/or detected. Here, concentration changes and not structural changes of the particles/molecules to be characterized in accordance with the inventive methods are measured/detected. Structural changes in this context are related to thermal denaturation mentioned in the first embodiment. The second illustrative embodiment illustrates that conformational changes and changes in surface (like size and chemistry) and interactions may be measured by thermo-optical characterization because the thermophoretic properties are altered. Also, thermo-optical “trapping” devices are illustrative for this embodiment. Corresponding illustrations of the usefulness of this embodiment of the invention are also illustrated in the appended examples, e.g. the determination of hydrodynamic radius and interaction between proteins, the detection of interactions between biomolecules and discrimination of nucleic acids by size, the detection of binding of molecules to particles, the investigation of conformation, structure and surface of (bio)molecules, the detection of conformational changes, like folding/unfolding of biomolecules, the trapping of particles (e.g. the trapping of vesicular structures or lipids) or (bio)molecules, and the detection of covalent and non-covalent modifications of particles.

It is documented and exemplified herein below that, e.g. the thermophoresis of nucleic acids (in particular of DNA) is length/size dependent and the means and methods provided herein allow for the determination and elucidation of single versus double stranded DNA as well as the determination also of small nucleic acids up to e.g. 100, 300, 1000, or 5000 nucleotides or base pairs. A non-limiting example is illustrated in appended FIG. 5, wherein mobility in a temperature gradient is measured by the means and methods provided herein. Here it is shown that in particular the second embodiment of this invention allows for the distinguishing verification between length/size (in the particular example 20 mer versus 50 mer) and/or “strandness” of nucleic acid molecules (in the particular example single stranded DNA versus double-stranded DNA). Again, also this second illustrative embodiment of the present invention is not-limited to the detection of short DNA or the determination of double- or single-stranded nucleic acid molecules. Also interactions between particles/molecules, conformations, hydrodynamic radii, binding kinetics and stabilities of particles/molecules, e.g. proteins, nucleic acids (e.g. DNA, RNA, PNA, LNA), nanoparticles, beads, particularly microbeads, lipids, liposomes, vesicles, cells, biopolymers (hyaluronic acid, alginate and alike), two-dimensional lipid sheets, inorganic substances (e.g. carbon-nanotubes, buckyballs, etc), Poly-ethylenglycol (PEG) may be measured. The molecules mentioned above show e.g. differences in temperature stability. Illustrative molecule specific temperature ranges for the measurement of respective thermo-optical properties are given in table 1.

The examples for uses of the means, methods and devices disclosed herein are not to be considered to be limiting and illustrate the invention. In particular, the present invention and its corresponding means and methods are not limited to the use of the detection, measurement and/or verification of biomolecules, like nucleic acids or proteins/proteinaceous structures. As is evident from the invention as disclosed herein, also any temperature-sensitive system can be adopted to the methods and devices disclosed herein.

It is, e.g. feasible to measure also chemical reactions, like inorganic or organic reactions.

The person skilled in the art is aware that the invention as disclosed herein is only restricted by the fact that the reaction to be measured, detected, verified and or assessed has to take place in an solution that can be heated, in particular heated optically.

In some embodiments the device according to the present invention is based on a fluorescence microscopy setup with an excitation means, e.g. a Light Emitting Diode (LED) for excitation, an excitation/emission filterset, a specimen holder for a microfluidic chamber and a fast CCD Camera for spatial resolved recording of the fluorescence intensity. Such a fluorescence microscopy setup is well established in life science and other areas. According to the present invention, such a common setup is extended by an infrared (IR) laser whose radiation is focused. The laser may be arranged below the specimen holder such that the radiation is focused from below the specimen holder into the microfluidic chamber by an IR corrected lens (as illustrated in the appended figures, particularly FIG. 1). However, the laser, the detection means and the excitation means may be arranged on one common side of the specimen holder, e.g. below the specimen holder as depicted for example in FIG. 2. In one embodiment, the specimen holder is attached to the objective. Such a setup avoids relative movements of the specimen holder with respect to the objective. According to a further embodiment, it is possible to move the laser freely in the object plane by using two voltage driven infrared mirrors. It is further advantageous to use thin liquid films (approx. 1 μm to 500 μm, preferably 1 μm to 50 μm, more preferably 1 μm to 20 μm, even more preferably 1 μm to 10 μm), e.g. in thin liquid chambers, of biomolecule solutions with local coherent IR radiation. However, the method of the present invention is not limited to thin liquid chambers. An extension to μl-drops or nl-drops of aqueous solutions, capillaries and micro-well plates is possible as illustrated in the appended figures, e.g. FIGS. 2 and 16 to 24. According to a further embodiment, the infrared heating and fluorescence detection are realised through the same objective, which makes the setup much more flexible and compact (see e.g. FIGS. 2 and 16 to 24). The use of one objective to focus both electromagnetic radiation from the infrared and visible part of the spectrum comprises the necessity that the objective does both with high optical quality. In particular the infrared radiation must not exhibit a strong dispersion by the objective. Strong dispersion would lead to a high temperature offset and a comparably strong temperature gradient at distances from the heated centre. This situation is in some embodiments described here avoided. Strong dispersion increases the time of measurement and decreases the precision. Without being limited by theory, this may be due to an increase of the length scale where thermophoresis is strong whereby the system needs more time to reach steady state. The second effect may be due to the fact that the nonlinear bleaching correction is only precise if thermophoresis is negligible at greater distance to the heat spot. This is achieved when diffraction of the infrared radiation is low. Accordingly and in accordance with this invention, only one direction in space may be used for detection and manipulation. The described method and the herein disclosed devices may be integrated into established instruments and high throughput systems.

Water shows a strong absorption of radiation in the infrared regime larger than 1200 nm. The absorbed energy is converted into heat. Coherent IR LASER and IR optics allow the creation of a very high power density of infrared radiation in solution. By controlling the LASER optics the LASER focus can be moved and changed. This includes optics which change the aspect ratio of the radial symmetric laser beam to produce a line shaped focus. This is in particular useful if measurements are conducted in a capillary. Since the whole cross section is heated homogeneously, a spatial temperature profile exist only along the length of the capillary. A temperature gradient in only one direction of space increases the precision of the measurement since the all pixels with same distance from the heated centre can be averaged. In particular this allows the use of a CCD camera with a single line of pixels. In this case the integration of fluorescence is obtained by hardware. By using a photodiode or photomultiplier the fluorescence from a finite volume element (i.e. from the centre of the heat spot/line) is measured, without any spatial resolution. A spatial resolution of fluorescence detection is only necessary in cases were the hydrodynamic radius is the thermo-optic property of interest. The optical heating technique allows the creation of broad temperature distributions and strong temperature gradients on the micron scale. At the position of the LASER focus there is the highest temperature. This upper temperature limit may be adjusted by controlling the power of the LASER and the shape of the laser focus. With increasing distance to the laser focus the temperature of an aqueous solution is decreasing due to thermal conductivity. The lower limit of the temperature may be set by the temperature of the surrounding chamber material. This material can be cooled down, e.g. up to 0° C. In this way it is possible to generate a temperature distribution containing all temperatures between 100° C. (with high laser power) at the laser focus and 0° C. at greater distances to the heat spot.

With the method of the present invention it is possible to heat and analyze solutions, particularly aqueous solutions in a thermo-optical way. There is no need for heat conducting materials like heat transducers from a heating element (copper wire, Peltiers etc.). The solution itself is directly heated by the LASER light. Because the laser focussing is only diffraction limited, temperature distributions spanning all temperatures between 0° C. and 100° C. (the complete liquid phase of water) can be observed simultaneously on a length scale of a few hundred micrometers.

With the method of the present invention, measurements are 3000-10000 times faster than the fastest available measurement systems known in the art. The method of the present invention allows to obtain all temperatures between 0° C. and 100° C. at the same time because a spatial temperature distribution is used. The temperature is not created by contact with a heating element, but within the sample itself. By using the infrared scanning optics, arbitrary two dimensional temperature patterns can be created in solution. This way, any structuring of the surface is obsolete. In addition, all materials transparent for radiation in the infrared can be used to build a microfluidic measurement chamber (glass, sapphire, plastic, silicon, crystals).

In addition, the method of the present invention can also be used to create temperature distributions in the aqueous solution near a surface. Because of the continuity of temperature the surface also adopts a temperature distribution. Therefore it is possible to heat the surface as well as the solution. A possible application is the analysis of DNA-microarrays. Temperature gradients close to a surface may be used to move molecules toward a surface or away from a surface. These local concentration changes can be measured precisely by total internal reflection fluorescence (TIRF) system shown in the appended figures, particularly FIGS. 24 and 36, or any optical system (1) capable of TIRF. This thermophoretic motion in the direction of the incident laser light can be used to direct molecules into a microfluic structure in order to capture and/or concentrate them. Therefore, temperature gradients generated according to the present invention can also be used to capture and/or concentrate molecules/particles. The capture and concentration of molecules/particles is dependent on their thermo-optical properties (e.g. the sign of the thermophoretic effect).

One way to suppress secondary effects related to inhomogeneous temperatures in solution can be suppressed by choosing the right microfluidic chamber geometry. For example convection is dealt with by using only a thin sheet of liquid. This also means that it is advantageous for reproducible and precise measurements that the height of the thin sheet of liquid does not vary from measurement to measurement. The speed of the convective flow, which is accompanied by the spatial temperature distribution depends quadratically on the height of the liquid sheet. This nonlinearity means that slight changes in chamber height lead to comparatively strong changes in the speed of the convective flow which in turn effects the concentration and temperature profile in a very complicated way. Therefore experiments are preferably performed in microfluidic measurement chambers of defined height (e.g. capillaries). Since in accordance with the means and methods of this invention temperature is generated herein by the generation of heat due to the absorption process, the constant height is also advantageous to obtain reproducible temperature distributions. Differences in height would lead to deviations due to differences in the amount of energy absorbed and because of differences in the volume/surface ratio. This ratio determines the rate of heat transfer to the surrounding and therefore also the temperature distribution in solution. The reproducibility of the temperature profile determines the maximal possible measurement precision.

Another way is to measure even faster to avoid disturbances by convection opening the possibility to measure in single droplets or micro-well plates (thicker sheet of liquid). Since the IR laser is absorbed on a length scale of 300 μm (1/e) thin samples, e.g., thin chambers are heated homogeneously in the z-direction (height).

Accordingly, the present invention provides an improved method to measure thermo-optically characteristics of particles/molecules in a solution with the steps of (a) providing a sample probe with marked particles/molecules in a solution; (b) exciting fluorescently said marked particles and firstly detecting fluorescence of said excited particles/molecules; (c) irradiating a laser light beam into the solution to obtain a spatial temperature distribution in the solution around the irradiated laser light beam; (d) detecting secondly a fluorescence of the particles/molecules in the solution at a predetermined time after irradiation of the laser into the solution has been started, and characterizing the particles/molecules based on said two detections.

In one embodiment the predetermined time is within the range of from 1 ms to 250 ms. Preferably the detection time is in the range of from 1 ms to 50 ms. In a particular embodiment, the laser beam is defocused such that a temperature gradient within the temperature distribution is in the range of from 0.0 to 2K/μm, preferably from 0.0 to 5K/μm. Preferably the laser beam is irradiated through an optical element into the solution. In a particularly embodiment the optical element is a single lens. In a particular embodiment of the invention, the method is further comprising the step of measuring the temperature distribution in the solution around the irradiated beam with a temperature sensitive dye. The temperature distribution may be determined based on detected fluorescence of the temperature sensitive dye, wherein the solution comprising said temperature sensitive dye is heated by the irradiated laser beam and the fluorescence spatial fluorescence intensity is measured substantially perpendicular around the laser beam. In a further embodiment the predetermined time is within the range of from 0.5 s to 250 s. Preferably in said predetermined time concentration change(s) within the spatial temperature distribution in the solution due to thermophoretic effects and such (an) concentration change(s) is(are) detected by a change of the distribution of fluorescence. In some embodiments the laser beam is focused such that a temperature gradient within the temperature distribution is achieved in the range of from 0.001 to 10K/μm. In a further embodiment of the invention fluorescence is detected with a CCD camera. In some embodiments, the brightness of said fluorescence is detected with a photodiode or a single pixel with the CCD in the centre of the laser beam. In further embodiments the particles are biomolecules and/or nanoparticles and/or microbeads and/or combinations thereof. In particular embodiments, the laser light is within the range of from 1200 nm to 2000 nm. Preferably the laser is a high power laser within the range of from 0.1W to 10W, more preferably of from 0.1W to 10W, even more preferably from 4W to 6W. In some embodiments, the solution is an aqueous solution with an particle concentration within the range of from 1 atto Molar (single Particle Microbeads) to 1M, preferably from 1 atto Molar to 100 μMolar. Particularly, the solution is a saline solution with concentrations in the range of from 0 to 1M. Preferably, the spatial temperature distribution is between 0.1° C. and 100° C. In preferred embodiments, the temperature gradient is created within 0.1 μm to 500 μm in diameter around the laser beam. The irradiation of the laser and the detection of the fluorescence is in a preferred embodiment of the invention conducted from the same side with respect to the sample probe. Preferably, the solution is provided with a thickness in direction of the laser light beam from 1 μm to 500 μm. In particular embodiments, the fluorescence is detected within a range of from 1 nm to 500 μm, particularly from 50 nm to 500 μm in direction of the laser beam. Preferably, the fluorescence is detected substantially perpendicular with respect to the laser light beam with a CCD camera. More preferably the second fluorescence detection is spatial measurement of the fluorescence in dependence of the temperature distribution substantially perpendicular with respect to the laser light beam. In preferred embodiments, the sample solution is in a capillary.

The present invention also provides a device for measuring thermo-optically characteristics of particles in a solution as described in any of the above embodiments, wherein the device comprises: a receiving means for receiving marked particles within a solution; means for fluorescently exciting the marked particles; means for detecting the excited fluorescence in said solution; a laser for irradiating a laser light beam into the solution to obtain a spatial temperature distribution in the solution around the irradiated laser light beam. In some embodiments, the means for fluorescently exciting the marked particles is a LED. Preferably, the laser is a high power laser within the range of from 0.1W to 10W, preferably 1W to 10W, more preferably from 4W to 6W. In a particular embodiment, the laser and the means for detecting the excited fluorescence are arranged on the same side with respect to the receiving means. In a more preferred embodiment, the device further comprises an optic for magnifying the detected region. In particular embodiments, the device further comprises an optic for focusing or defocusing the laser beam. Preferably the optic is a single lens. In preferred embodiments the detecting means is a CCD camera. In some preferred embodiment, the CCD camera is a line CCD camera. In particular embodiments, the detection is one-dimensional along the length of a capillary. In another particular embodiment, the detecting means is a photo diode.

The present invention also relates to the use of the methods and the devices described in any of the above embodiments for detecting and/or measuring the characteristics of particles and/or molecules in solution. The molecules to be detected, measured or characterized in accordance with this invention may also be drug candidates.

In particular embodiments of the invention, the characteristics to be detected or measured in accordance with this invention are selected from the group of stability, length, size, conformation, charge, interaction, complex formation and chemical modification of particles. In preferred embodiments, the particles to be measured are selected from the group consisting of (a) molecule, biomolecule(s), nanoparticles, beads, microbeads, (an) organic substance(s), (an) inorganic substance(s) and/or combinations of these. Preferably said particle is selected from the group consisting of (a) (bio)molecule(s), nanoparticles, microparticles, microbeads, (an) organic substance(s), (an) inorganic substance(s) and/or combinations of these. More preferably, the (bio)molecule is selected from the group consisting of (a) protein(s), (a) peptide(s), (a) nucleic acid(s) (e.g. RNA (e.g. mRNA, tRNA, rRNA, snRNA, siRNA, miRNA), DNA), (an) RNA aptamer(s), (an) antibody/antibodies (or fragments or derivatives thereof), (a) protein-nucleic acid fusion molecule(s), (a) PNA(s), (a) locked DNA(s) (LNAs) and (a) biopolymer(s) (sugar polymer, hyaluronic acids, alginate, etc.). Also intra- or inter-molecular interactions, e.g. protein folding/unfolding, are within the scope of this embodiment.

A particularly preferred embodiment of the present invention relates to a method to measure thermo-optically the physical, chemical or biological characteristics of particles/molecules in a solution with the steps of (a) providing a sample probe with marked particles/molecules in a solution in a capillary; (b) exciting fluorescently said marked particles/molecules and firstly detecting fluorescence of said excited particles/molecules one-dimensionally along the length of the capillary; (c) irradiating a laser light beam into the solution to obtain a linear temperature distribution in the solution around the irradiated laser light beam along the length of the capillary; and (d) detecting secondly a fluorescence of the particles/molecules in the solution at a predetermined time after irradiation of the laser into the solution has been started, and characterizing the particles based on said two detections.

The present invention also provides a particular device for measuring thermo-optically the physical, chemical or biological characteristics of particles/molecules in a solution as described in any of the above embodiments, wherein the device comprises: a capillary for receiving marked particles/molecules within a solution; means for fluorescently exciting the marked particles/molecules; means for detecting the excited fluorescence in said solution one-dimensionally along the length of said capillary; a laser for irradiating a laser light beam into the solution to obtain a linear temperature distribution in the solution around the irradiated laser light beam.

The device according to the present invention with a fluorescence microscopy and local infrared laser heating may also be used to measure the effect of temperature gradients (i.e. temperature inhomogeneities) on dissolved molecules (see second embodiment of the present invention). Almost all dissolved molecules start to move in a temperature gradient, either to hot or cold regions. This effect is called thermophoresis or Soret effect and is known for 150 years. But the mechanism of molecule movement in liquids stayed unclear. A mayor step towards a theoretical understanding of thermophoresis in liquids has been done recently.

With the method and the device according to the invention, the stability, conformation, size and/or length of molecules, in particular biomolecules may be characterized and/or determined. The interaction of (bio)molecules with other molecules or particles, e.g. further (bio)molecules, nanoparticles or beads, e.g. microbeads is characterized in particular embodiments of the invention. The molecules to be analyzed may also be linked (e.g. covalently or non-covalently) to beads or particles, e.g. the beads or particles may be coated with molecules (e.g. biomolecules) to be analyzed, characterized in accordance with this invention.

In the following two illustrative methods according to the present invention which rely on a very similar measurement protocol but analyze very different molecule parameters will be discussed. Only in the method of the second illustrative embodiment of the present invention the thermophoretic motion of particles is used. In the method of the first illustrative embodiment this effect has to be excluded. Furthermore, particular embodiments of the present invention will be explained in detail with reference to the figures and with reference to the appended detailed examples. Said references to the figures and examples are not considered to be limiting.

First Illustrative Embodiment of the Invention

The method according to a first illustrative embodiment is in particular useful in a temperature stability measurement of molecules, in particular biomolecules. However, it is again of note that the means and methods provided herein are not limited to the detection, verification and/or measurement of biomolecules. The method described and illustrated as a non-limiting example, in the following allows, e.g., the measurement of melting temperatures (FIG. 3) (stability, thermodynamic parameters like dS (change in entropy), dH (change in enthalpy) and dG (change in Gibbs free energy)) of biomolecules (proteins, double stranded (ds)RNA, dsDNA were one nucleic acid strand could also be bound to a (nano)particle, microbead, surface etc). With said method measurements of melting curves of dsDNA and DNA hairpins have been conducted. The results are very well comparable to respective literature values. As mentioned herein above, the present invention is in particular useful in the measurement of biomolecules in general. Illustratively, it is shown hereon that e.g. SNPs (single nucleotide polymorphisms) in (short) DNA strands can easily be detected (see also FIG. 4).

A particular embodiment of the first illustrative method will be explained in the following. Nucleic acids with a fluorescence tag are given into a thin microfluidic chamber (i.e. e.g. 40 μm, 20 μm, 10 μm, or 5 μm, preferably 20 μm). The modification of nucleic acids with tags, like fluorescent tags is a well established technique which is broadly used. Before heating starts, the fluorescence is observed to determine the fluorescence level of 100% not melted molecules. It is important that the used fluorescence tag reacts on the melting of the two DNA strands (or RNA strands or protein structure, or the fluorescence of a nanoparticle reacts on whether ssDNA/RNA or dsDNA/RNA are bound to it). This can be realized by using e.g. fluorophore/quencher pair (Donor/Quencher pair particularly Donor/Acceptor pair: Energy Transfer (ET), e.g. Resonance Energy Transfer (RET), particularly Fluorescence Resonance Energy Transfer (FRET)) or by the dissociation of an intercalating nucleic acid stain (e.g. SYBR Green/POPO/YOYO) or a protein stain (e.g. SYPRO Orange (Invitrogen)). In case of e.g. gold-nanoparticles, fluorescence changes by changing of index of diffraction by DNA binding. The laser coupled to the microscope is defocused and adjusted in a way that the temperature gradients are low enough to decrease the thermophoretic particle drift to negligible values. At the same time focusing must be tight enough to reach temperature high enough for melting of the molecules. The measurements is in some particular embodiments performed with high temporal resolution in the microsecond range, since the measurement has to be performed in a time span were thermophoretic motion is still negligible but the heating process of the microfluidic chamber is completed. The measurement is strongly dependent on the experimental conditions. The data necessary for determining the melting temperature are typically obtained within 200 ms, within 150 ms, within 100 ms or within 50 ms, preferably within 150 ms. This is very surprising and supports the gist of this invention. Even shorter time spans are possible. In context of this invention, e.g. for a qualitative discrimination of different species or nucleic acids. A first image is taken before the IR laser is turned on to obtain the fluorescence level of 100% not melted molecules. A second image is taken 100 ms, preferably 50 ms, more preferably 40 ms after the laser is turned on (were the chamber has reached its steady state temperature). These two images contain all necessary information. The first image contains the fluorescence of the not melted molecule. The image taken while the laser is turned on allows the observance of the percentage of melted molecules at all different temperatures simultaneously, ranging from 0% far away from the heat spot (cold) to 100% in the centre of the heat spot (hot). From an independent measurement the temperatures of all pixels in the melting experiment are known. Plotting the percentage of i.e. melted DNA strands vs. temperature allows to determine the stability of the molecule (i.e. the melting temperature) and to derive the thermodynamic parameters.

In summary, the first illustrative embodiment of the present invention connects the measurement of a spatial temperature distribution on the micrometer scale with the measurement of temperature dependent chemical/biochemical reactions and temperature dependent interactions between molecules/biomolecules/nanocrystals/microbeads. For the measurements the absolute temperatures not the temperature gradients are important. One can consider each detection volume (mapped onto a CCD-camera-pixel) as an microreactor. For the measurement it is very important that every molecule stays in this microreactor during the time of measurement. Therefore the measurement has to be as fast as to prevent thermophoresis and convection to move the particle out of the area addressed as microreactor. This requirement is satisfied with measurement time of 150 ms.

With the method and the device according to the first illustrative embodiment one is able to discriminate between dsDNA molecules with a single nucleotide mismatch (SNP, single nucleotide polymorphism), as well as salt dependent and length dependent differences in stability. The differences in stabilities can also be measured for dsRNA, DNA and RNA hairpins (ssDNA/RNA base pairing with themselves) and proteins. Also the influence of different aqueous buffers systems to the stability can be measured (pH, salt concentration, valency of ions). The biomolecules can also be coupled to (nano)particles, (micro)beads. The fluorescence of these modified particles changes depending on whether e.g. ssDNA/RNA or dsDNA/RNA (or other biomolecules) is bound to it. Heating of such solutions leads to the same result without the need of specific fluorescent markers bound to the biomolecule. Beside a qualitative discrimination, also a quantitative thermodynamic analysis is possible. Since the measurement times are in some cases below the relaxation times of intermolecular reactions, it is not in all cases possible to determine the thermodynamic parameters like dS (change of entropy), dH (change of enthalpy) and dG (change of Gibbs free energy) directly. But by using non equilibrium thermodynamics, these can be easily calculated. For the qualitatively discrimination of (i.e.) single nucleotide polymorphisms, working under non equilibrium condition can increase the measured differences in stability of the compared molecules. The measurement of mismatches in nucleotide sequences is of great importance in medical diagnostics. The method allows identifying hereditary diseases. It can also be used in pharmaceutical high throughput screenings for the binding of low molecular weight compounds to nucleic acids. Furthermore the melting of double-stranded (ds) nucleic acids allows to determine their length.

The measured melting curves of the present invention reproduce the results measured by established techniques but up to 3000× faster than Peltier or heating bath based methods, like PCR cycler or fluorimeter methods. The inventive method is much faster since it is not necessary to heat the volume by direct contact, and the reaction of molecules on a specific temperature between 0° C. and 100° C. is observed at the same time. Again, the gist of this embodiment is that temperatures are generated and measured with spatial instead of temporal resolution. There are no delays due to heating and cooling times which makes the method of the present invention very fast. At the same time only thin sheet of liquids are used which decreases the necessary sample volume. In addition the manipulation and analysis of the molecules occurs all optically without the risk of contamination. This is essential if the analysis is combined with PCR reactions where any contamination with i.e. human DNA renders an analysis impossible.

By performing the thermal denaturation within e.g. 100 ms, preferably 50 ms, it is possible to measure the effect of substances on the DNA/protein stability which are sensitive to high temperatures (e.g. DNA binding proteins, substances like cyclic-Adenosin Monophosphate (cAMP)). These substances will be damaged or degraded in an experiment with techniques of the prior art and an effect on the thermal stability is not detectable with such prior art techniques.

Within the spatial temperature distribution there are temperature gradients. If the measurement time is longer than the 150 ms (for the here proposed chamber thickness of e.g. 20 μm) one can measure the thermophoretic movement of the biomolecules (molecules, nanoparticles, microbeads). From this measurements additional information can be gathered.

Second Illustrative Embodiment of the Invention

The method according to a second illustrative embodiment (i.e. the above recited method relating to measurements wherein the pre-determined time in said second detection of the method of the invention is within a range of 0.5 seconds to 250 seconds, preferably within a range of 0.5 s to 50 s, more preferably within a range of 0.5 s to 40 s) is in particular useful for a measure of the mobility of molecules in a temperature gradient and its use for biomolecule characterisation. The method of the first embodiment described above analyzes molecules in a temperature distribution on a short timescale of milliseconds. Dynamics effects like thermophoresis can be neglected in this short time interval. If molecules are observed for a time period in the order of seconds, thermophoresis sets in and molecules start moving in the temperature gradient. This effect drives molecules analogous to electrophoresis along a gradient to lower temperatures (in some cases the opposite is also observed). The velocity of the molecules is directly proportional to the temperature gradient with a molecule specific coefficient D_(T) (thermophoretic mobility): v=−D_(T)∇T.

Unexpectedly the thermophoretic mobilities of biopolymers vary strongly with the chain/molecule length.

The thermophoretic mobility of these molecules varies strongly with molecule parameters which change the entropy of solvation, size, charge, kind of surface, size of surface, hydrodynamic radius etc. This opens the possibility to discriminate biomolecules and detect an interaction between them (also between nanoparticles/micro beads and biomolecules) (as illustrated in the appended figures, particularly FIG. 4).

Since thermophoresis builds up concentration gradients, the effect is counteracted by ordinary diffusion. The interplay between these two effects leads to a steady state concentration profile which is expressed by the following equation

$\frac{c}{c_{0}} = {{\exp\left\lbrack {{- S_{T}} \times \Delta\; T} \right\rbrack}.}$ The concentration at any given point in a temperature distribution is solely dependent on the difference in temperature and not the temperature gradient any more. The quotient of thermophoretic mobility D_(T) and ordinary diffusion constant D is called Soret coefficient S_(T) and describes the magnitude of thermophoresis in steady state. It is exponentially dependent on the temperature difference. Thus the precision of the measurement is strongly dependent on the reproducibility of the temperature profile.

A typical measurement procedure according to the second embodiment will be described in the following. In the very beginning an image is taken without IR LASER heating to determine the fluorescence intensity of the 100% relative concentration level. Then the LASER is turned on. In this experimental setup it is possible to focus the laser tightly with a half width below 6 μm to create strong temperature gradients or to use it defocused (as described above with a temperature profile half width of e.g. 200 μm). This influences the velocity of the molecules and how fast the steady state is reached. The necessary temperature increase varies between 0.1° C. and 80° C. above ambient temperature (20° C.). If the chamber is e.g. cooled to 0° C., a temperature range between 0.1° C. and 100° C. temperature increase can be realized, dependent on the thermal stability of the sample and the magnitude of the thermophoretic effect. In general no high temporal resolution of the image recording is necessary to determine the thermophoretic effect of the (bio)molecules, (nano)particles, or (micro)beads. The measured signal is the steady state or close to steady state concentration which is in most cases reached after a few seconds. The signals of most molecules differ strong enough from each other before steady state is reached to identify them without any doubt. For the data analysis beside the initial concentration the concentration at a certain time after LASER heating is turned on is needed (also time courses can be taken). It is sufficient to determine the concentration of a single pixel (i.e. point of maximum temperature).

Before the laser is turned on there is a homogenous distribution of the biomolecules. Therefore the fluorescence intensity (or whatever the measurement signal is), which is direct proportional to the concentration, has the same magnitude at each point. When the LASER is turned on the concentration distribution changes. The molecules are moving away from the hot laser focus. Therefore the magnitude of the fluorescence intensity is decreasing until the steady state is reached. This decrease can be measured and thus the characteristics of the molecules can be derived by the theory of thermophoresis and different molecules can be discriminated by comparison.

With the methods as disclosed herein, in particular in relation to the second embodiment, for means and methods are provided to detect, measure and/or verify a large variety of interactions. For example DNA/DNA, RNA/RNA, protein/protein, protein/DNA, protein/RNA interactions, but also protein, DNA, RNA interaction with other materials like nanoparticle/microbeads can be measured. The only requirement is a label, in particular a fluorescent label, bound to one of the molecules. An exception is, e.g. the case of large microbeads, where light scattering can be used (directly) for detection. If one uses modified microbeads in a manner that for example DNA single strands can bind to this modifications, the mobility of this bead is changed due to the binding. Hence, the method of the second embodiment allows to detect this binding. Because this modified beads are used in sequencing setups the method of the present invention can be applied there. With the present method it is possible to detect everything that changes the size, charge or surface of a molecule. It has been shown that the method of the present invention is also able to measure specific DNA binding to nanoparticles and polystyrene microbeads via Streptavidin/Biotin (see FIG. 35). Also interactions between antibody and an epitope are detectable. Also the binding of a protein to a DNA strand, for example of a polymerase, is detectable.

Since no high temporal and spatial resolution is needed the method of the second embodiment is cost-efficient and easy to realize. For example instead of a CCD camera an avalanche photodiode can be used (only information of a single Pixel is needed). If the microfluidic chamber is a capillary (i.e. a microfluidic chamber with high aspect ratio (Length/Width)), no spatial resolution in direction of the width is needed and for the detection of fluorescence distribution only a line-CCD-camera is needed. This alternative in between CCD-Camera and avalanche photodiode/photomultiplier is very cost efficient and safes time for data evaluation since the integration of fluorescence is performed by the hardware. The measured systems can exhibit concentration down to nanomolar without any restriction at higher concentrations. Also a comparable high degree of contamination is tolerated by the methods of the present invention. Measurements are also possible in crude extract of cells or in blood. Beside tolerance to contaminations the method also sustains strong variations in the viscosity of the solution. Measurement can for example be performed in water or glycerol or in aqueous solution with a gel like consistence. Since measurements are performed in microfluidic chambers the volume needed for an experiment is e.g. only 0.5 μl, 1 μl, 2 μl, 5 μl, 10 μl, preferably 2 μl, and can be further reduced. Because of its easy calibration the method has a big advantage compared to FCS (Fluorescence Correlation Spectroscopy) and can be easily automated. The method is distinguishable faster than any other method on the market to determine interactions between biomolecules/nanoparticles/microbeads (i.e. Biacore). The length of short DNA molecules is determined within seconds compared to an hour which is needed by gel electrophoresis, an established method in this field. A further advantage is that the measurements are performed in aqueous solution. It is not necessary to change the phase in which the molecule is dissolved (gel in gel electrophoresis or C18, HIC-columns in HPLC). The possibility to differentiate between single and double stranded DNA that fast opens up new possibilities in diagnosis as well as scientific research. One example is, e.g. the diagnosis of infectious diseases, like viral diseases or bacterial infections.

In summary, the inventive method of the second embodiment manipulates concentrations of biomolecules/nanoparticles/microbeads in aqueous solutions by temperature gradients (up to 10K/μm, in some embodiments up to 5K/μm, particularly up to 2K/μm) established with an IR LASER. In the enclosed example a general theory for thermophoresis in liquids is described.

The most important features and advantages of the present invention will be summarized in the following. The method and the device of the present invention works all optical, i.e. manipulation and detection is made by optical means. With the present invention, it is possible to optically manipulate molecules down to the size of a single fluorescent dye which is not possible by optical traps which are limited to spherical particles of 500 nm. The method is free of contaminations and easy to miniaturize and to parallelize. That makes it possible to integrate the system in established instruments like pipetting robots etc. Heating of aqueous solutions on the micron scale allows to create temperature distributions which renders long heating and cooling periods unnecessary. The criterions for materials which can be used to build measurement chambers are very unspecific. Another advantage is that IR LASER are getting more and more common in the telecommunication industry and are produced in big quantities. Furthermore the technique of fluorescent dye coupling to biomolecules has become a cost-efficient standard technology. Thus, with the present invention, it is possible to measure stability of molecules as well as any kind of interaction (with each other, the buffer system, other solutes, etc.). Therefore, the invention is not limited for use in the measurement, detection and/or verification of biomolecules or biological, biomedical, biophysical and/or pharmacological (in vitro) processes.

With the means and methods of the first embodiment of the present invention, melting curves over a large range of temperatures can be measured and determined. According to this embodiment, thermophoresis should be avoided. The temperature achieves a change in the structure, which is detectable via the fluorescence behaviour. The characteristics can be detected and/or measured over a wide temperate range at one time. None of the known prior art documents discloses an inter- and/or intramolecular reaction induced via temperature. According to the first embodiment, the laser can be irradiated into the solution via optical fibres. The light may exit the optical fibres divergent. According to a preferred embodiment of the present invention, the optical system for focussing the laser can be a single lens.

The means and methods of the second illustrative embodiment of the present invention provides the advantages that the effect of thermophoresis is used in a controlled manner. In particular, concentration changes, induced by thermophoresis effect(s) are measured via a change in fluorescence behaviour. Thus, the fluorescence signal recorded in the second embodiment is primary based on changes of concentration and not on changes in the structure of the tested particles or biomolecules. The second illustrative embodiment includes that changes in concentration are sensitive to changes in the structure of a particle or molecule. In the known prior art only a maximum temperature difference of only 2.5K is disclosed. Moreover, only measurements with a low power laser (320 mW) are described in the prior art. Particularly, a CCD camera can be used for detecting or measuring the fluorescence of the sample. According to another embodiment, only one pixel of the CCD camera has to be used for the measurement or the detection of the fluorescence light, e.g. only a central 1×1 μm pixel has to be used. This has the advantage, that further spatial information may be neglected according to certain embodiments of the present invention.

According to a further aspect of the present invention, temperature distributions around the irradiated laser light beam are measured with independent measurements. These measurements are typically based on known temperature-dependent fluorescence behaviour of dyes.

Accordingly, as also illustrated in the appended examples, particular embodiments of the present invention relate to the detection of thermodiffusion or thermophoresis of (bio)molecules or particles, the determination of hydrodynamic radii (bio)molecules or particles, the detection of binding of or between (bio)molecules or particles, the detection of interactions of or between (bio)molecules or particles, the detection of conformational changes in (bio)molecules, the detection of denaturation of proteins or melting of nucleic acids and to optothermal trapping of (bio)molecules or particles.

BRIEF DESCRIPTION OF THE DRAWINGS

In the following description of examples of preferred embodiments of the invention, elements having the comparable technical or physical effect have the same reference numerals.

FIG. 1A shows an fluorescence IR scanning microscope according to the present invention. The IR scanning microscope is based on a standard fluorescence microscopy setup (e.g. Zeiss AxioTech, Vario). Said device comprises: one or more light sources 32, preferably high power LED (e.g. V-Star, Luxeon) to excite the particles. The signal from the particles may be collected with an optical system 1, preferably a 40× oil objectiv and is seperated from the light of the light source by one or more light separation elements 4, preferably a dichroic mirror. The said signal is recorded with one or more detectors 31, preferably CCD Camera (e.g. SensiCam QE, PCO). The beam of an IR LASER 30 (e.g. IPG, Raman fibre RLD-5-1455) is coupled into microfluidic chamber 45 (e.g. a multiwell plate). The system may comprise other components as would ordinarily be found in fluorescence and wide field microscopes. Examples for means of excitation and detection of fluorescence may be found in: Lakowicz, J. R. Principles of Fluorescence Spectroscopy, Kluwer Academic/Plenum Publishers (1999).

FIG. 1B shows a further embodiment of an fluorescence IR scanning microscope according to the present invention similar to the embodiments shown in FIG. 1Aa. However, the light source 32 is oriented in a different way with respect to the dichroic mirror 4. The testing sample 50 is sandwiched between two pieces of glass 51, preferably coverslips.

FIG. 2 shows another embodiment of an IR scanning microscope according to the invention. According to this embodiment of the present invention, the testing sample 50 is provided in form of a single droplet with the marked particles. The volume (preferably some nanoliter up to some microliter) of such a droplet can be easily adjusted such that also the dimensions, i.e. the thickness of the droplet which is irradiated with the laser beam, is predictable. In this embodiment, the laserbeam, the fluorescently exciting light from the light source 32, preferably a LED as well as the measured fluorescent light are all focussed by a common optical system 1, preferably a microscope objective with high numerical aperture and more preferably a objectiv with high numerical aperture and high IR transmission. Therefore, the LED, the LASER 30 and the detector 31, preferably a CCD can be arragend at a common side with respect to the sample. The exciting light is separated from the fluorescent light by a light separation element 5, preferably a dichroic mirror, which preferably seperates different parts of the light spectrum than the light separation element 4.

FIG. 3A-D show how melting curves with an radiation of 150 ms can be taken. (A) By measuring the fluorescence of a temperature sensitive dye, the temperature distribution in the microfluidic chamber can be measured. (B) shows the radial average of the temperatures measured by fluorescence. (C) Aprox. 150 ms after the IR LASER is turned on an image of the fluorescently labelled DNA is taken. The high intensity shows melted ds DNA. From (A) and (C) the melting curve can be determined very fast (D).

FIG. 4 shows a fast SNP detection. A 16 mer of dsDNA with a single nucleotide mismatch in the center (blue) is compared with the wild type (black). Within 150 ms both species can be clearly discriminated. The ordinate describes the fraction of dissociated dsDNA. The melting point is defined by 50% dissociated molecules. It is shifted by 15° C. by a single mismatch.

FIG. 5A-5B show the mobility in a temperature gradient. The figures show the change of concentration in the central pixel of a heat spot over time. A few seconds after the start of the measurement the laser is turned on and the concentration decreases until a steady state is reached. The signal allows to distinguish between a 20 mer and 50 mer dsDNA (A) as well as between 20 bases ssDNA and 20 basepair dsDNA (B).

FIG. 6 is an example for a fluorescence dye to be used in the methods of the present invention (6-carboxy-2′,4,4′,5′,7,7′-hexachlorofluorescein (HEX, SE; C20091) from Invitrogen).

FIGS. 7-14 show further information with regard to a detailed example according to the present invention. These figures show in particular:

FIG. 7: illustrates how thermodiffusion manipulates the DNA concentration by small temperature differences within the bulk solution. A thin water film is heated by 2 K along the letters “DNA” with an infrared laser. For a cooled chamber at 3° C., fluorescently tagged DNA accumulates to the warm letters. However at room temperature DNA moves into the cold, showing reduced fluorescence. The chamber is 60 μm thin, containing 50 nM DNA in 1 mM TRIS buffer. Every 50th base pair is labelled with TOTO-1.

FIG. 8A-B illustrates the salt dependency. (A) Thermodiffusion in water is dominated by ionic shielding and water hydration. (B) Soret coefficient ST versus Debye length for carboxyl modified polystyrene beads of diameter 1.1, 0.5 and 0.2 μm. Linear plot (left) and logarithmic plot (right). The Soret coefficients are described by equation (2) with an effective surface charge of σeff=4500 e/μm2 known from electrophoresis. The intercept S_(T)(λ_(DH)=0) is fitted with an hydration entropy per particle surface of Shyd=−1400 J/(molKμm2).

FIG. 9A-B shows a temperature dependency. (A) The temperature dependence is dominated by the linear change in the hydration entropy Shyd. It shifts the salt dependent thermodiffusion ST(λDH) to lower values. The particle size is 1.1 μm. (B) The Soret coefficient ST increases linearly with the temperature as expected for a hydration entropy Shyd(T). It depends on the molecule species, not its size, as seen from the rescaled Soret coefficients for DNA with different lengths.

FIG. 10A-D shows a size dependency. (A) For polystyrene beads, the Soret coefficient scales with the particle surface over four orders of magnitude. Measurements are described by equation (2) with an effective surface charge density of σeff=4500 e/μm2 and negligible hydration entropy. The deviation for the bead with 20 nm diameter can be understood from an increased effective charge due to the onset of charge normalization for a≤λ_(DH). (B) Accordingly, the thermodiffusion coefficient DT scales linearly with bead diameter. (C) The Soret coefficient of DNA scales according to S_(T)∝√{square root over (L)} with the length L of the DNA based on equation (2) with an effective charge per base pair of 0.12 e. (D) Thermodiffusion coefficient DT decreases over DNA length with D_(T)∝L^(−0.25), caused by the scaling of diffusion coefficient D∝L^(−0.75).

FIG. 11A-B: shows an effective Charge from Thermodiffusion. Effective charge is inferred from thermodiffusion using equation (3). Polystyrene beads (20.2000 nm) (A) and DNA (50-50,000 bp) (B) are measured over a large size range, impossible with electrophoresis. As expected, the effective charge of the beads scales with particle surface and linearly with length of DNA.

FIG. 12A to 12D: shows the dependency of thermodiffusion from concentration over time. (A) Diffusion coefficient D was obtained from back diffusion after switching off the heat source. (B) D is varied until the finite element simulation matches the experiment. (C) Radial depletion of DNA from a focussed 2K heat spot is monitored over time. (D) Comparison with simulation with known D yields DT and ST.

FIG. 13 shows a scaling of DNA diffusion coefficients. The diffusion coefficients as measured in this study at room temperature. The scaling over DNA length matches literature values with two scaling regimes with exponent −1 for short and −0.6 for long DNA33. As approximation, diffusion across the two scaling regimes is well described with an overall exponent of −0.75.

FIG. 14A-B: shows a simulation of Microfluidic Heating. (A) A 10 μm thin water film is enclosed between PS walls. Low thermal conduction of the chamber walls allow a thickness independent temperature profile, confirmed by the shown finite element calculation. (B) Convection is slow at maximal velocities of 5 nm/s due to thin chamber and comparable broad heating focus.

FIG. 15 shows the temperature dependence of a fluorescent dye measured by a fluorimeter with temperature control (Peltier element).

FIG. 16 illustrates a particular embodiment of a device accordingly to the present invention. The device may have an substantially arbitrary orientation with respect to the direction of gravitation, preferably the device is oriented perpendicular with respect to the direction of gravitation, more preferably the device is oriented substantially parallel or anti-parallel with respect to the direction of gravitation. Preferably the orientation of the device with respect to the testing sample or chamber may be adjusted as shown in FIG. 1a , FIG. 1b and FIG. 2. The device comprises: 1: Objective (e.g. 40×, NA 1.3, oil immersion, ZEISS, “Fluar”); 20: Scanning module, may be Galvano scanning mirror or may be acoustic optic deflector (AOD); 3: cold mirror, preferably high IR-transmission and preferably >90% reflection 350 nm-650 nm; 11: beam shaping module to determine laser beam diameter and focusing, may be a lens system which may comprise one, two or more lenses; 16: Laser fibre coupler w/o collimator; 15: Laser fibre (single mode or multimode); 30: IR-Laser (e.g. 1455 nm, 1480 nm, 0.1W-10W); 4: Dichroic mirror/beam splitter reflecting short wavelength (R>80%), transmitting long wavelength (T>80%); 7: Emission Filter (band pass/long pass); 31: detector, may be a CCD-Camera, Line-Camera, Photomultiplier Tube (PMT), Avalanche Photodiode (APD), CMOS-Camera; 6: Excitation Filter (e.g. band pass/long pass); 10: lens system to determine the beam properties of the excitation light source, may comprise one, two or more lenses; 32: Excitation light source, may be Laser, Fibre Laser, Diode-Laser, LED, HXP, Halogen, LED-Array, HBO. Preferably the listed parts are enclosed in a housing; a cold mirror is a specialized dielectric mirror, a dichromatic interference filter that operates over a very wide temperature range to reflect the entire visible light spectrum while very efficiently transmitting infrared wavelengths.

FIG. 17 shows an embodiment of the invention according to FIG. 16, wherein the scanning module 20 is replaced by a fix mirror 21, preferably a silver mirror.

FIG. 18 shows an embodiment of the invention according to FIG. 16, wherein the scanning module 20 is replaced by a fix mirror 21, preferably a silver mirror, and wherein a shutter 33 for controlling the IR LASER radiation is added and wherein a line forming module 12, may be a cylinder lens system or preferably a Powell lens, are added.

FIG. 19 shows a further embodiment of the device according to the present invention, in particular a confocal setup. The device may have an arbitrary orientation with respect to the direction of gravitation, preferably the device is oriented substantially perpendicular with respect to the direction of gravitation, more preferably the device is oriented substantially parallel or anti-parallel with respect to the direction of gravitation. Preferably the orientation of the device with respect to the testing sample or chamber may be adjusted as shown in FIG. 1A, FIG. 1B and FIG. 2. The device comprises: 1: Objective (e.g. 40×, NA 1.3, oil immersion, ZEISS, “Fluar”); 2: hot mirror, high IR-reflection, visible light transmission>80%; 4: Dichroic mirror reflecting short wavelength (R>80%), transmitting long wavelength (T>80%); 7: Emission Filter (band pass/long pass); 31: Detector, may be Photomultiplier Tube (PMT), Avalanche Photodiode (APD); 13: Pinhole aperture; 10: lens system to determine the beam properties of the excitation light source, may comprise one, two or more lenses; 32: excitation light source preferably a laser, more preferably a fibre coupled laser 11: beam shaping module to determine laser beam diameter and focusing, preferably lens system which may comprise one, two or more lenses; 16: Laser fibre coupler w/o collimator; 15: Laser fibre (single mode or multimode); 30: IR-Laser (e.g. 1455 nm, 1480 nm, 0.1W-10W); 33: Shutter; 14: pinhole aperture in confocale position to laser pinhole 13 or to laser fibre coupler 17 whereas the pinhole 13 may not be needed, preferably if a fibre coupled laser is used as excitation light source 32; 18 Laser fibre may be single mode or may be multi mode.

FIG. 20 shows a further embodiment of the device according to the present invention. The device may have an arbitrary orientation with respect to the direction of gravitation, preferably the device is oriented substantially perpendicular with respect to the direction of gravitation, more preferably the device is oriented substantially parallel or anti-parallel with respect to the direction of gravitation. Preferably the orientation of the device with respect to the testing sample or chamber may be adjusted as shown in FIG. 1A, FIG. 1B and FIG. 2. The device comprises: 1: Objective (e.g. 40×, NA 1.3, oil immersion, ZEISS, “Fluar”); 2: hot mirror, high IR-reflection, visible light transmission>80%; 4: Dichroic mirror reflecting short wavelength (R>80%), transmitting long wavelength (T>80%); 6: Excitation Filter (e.g. band pass/long pass); 7: Emission Filter (band pass/long pass); 10: lens system to determine the beam properties of the excitation light source; 31: Detector, may be Photomultiplier Tube (PMT), Avalanche Photodiode (APD); 32: excitation light source 16: Laser fibre coupler w/o collimator; 15: Laser fibre (single mode or multimode); 30: IR-Laser (e.g. 1455 nm, 1480 nm, 0.1W-10W); a hot mirror is a specialized dielectric mirror, a dichromatic interference filter often employed to protect optical systems by reflecting heat back into the light source. Hot mirrors can be designed to be inserted into the optical system at an incidence angle varying between zero and 45 degrees, and are useful in a variety of applications where heat build-up can damage components or adversely affect spectral characteristics of the illumination source. By transmitting visible light wavelengths while reflecting infrared, hot mirrors can also serve as dichromatic beam splitters for specialized applications in fluorescence microscopy.

FIG. 21 shows an embodiment of the invention according to FIG. 20, wherein a shutter 33 is added to control the IR laser radiation.

FIG. 22 shows an embodiment of the invention according to FIG. 20, wherein a shutter 33 and a line forming module 12, preferably a lens system comprising one, two or more lenses or more preferably a Powell lens are added to control the IR laser radiation.

FIG. 23 shows an embodiment of the invention according to FIG. 16, wherein the scanning module 20 is replaced by a fix mirror 21, preferably a silver mirror and wherein an additional light separation element 5, preferably a dichroic mirror, which preferably separates different parts of the light spectrum than the light separation element 4 is added and wherein an emission filter 8 is added, preferably transmitting another range of wavelength than the emission filter 7 and wherein a second detector 31 is added which detects the signal passing through the filter 8.

FIG. 24 shows a further embodiment of the device; 31: CCD camera; 7: Emission filter (e.g. bandpass or longpass); 4: light separation element, preferably a dichroic mirror for splitting excitation light pathway and emission light pathway; 1: Objective; 45: Chamber; 10: Lens system for focusing IR-Laser onto sample; 20: scanning module, may be galvanometric mirrors or an acoustic optical deflector; 16: Laser fibre coupler with collimating optics; 15: Laser fibre; 30: IR-Laser; 6: Excitation Filter; 10: optical system for excitation, may be a lens system comprising one, two or more lenses; 32: Light source (HXP,LED); 46: xyz-Translation stage for laser positioning, may be automated, preferably may be automated for scanning the chamber. 47: optical system for beam shaping of the IR Laser, may be a lens system comprising one, two or more lenses or may be a objective, preferably with high IR transmission.

FIG. 25A-B shows a quantification of interaction between biomolecules. 100 nM of a fluorescently labelled antibody (anti-Interleukin 4) are titrated with various amounts interleukin. (left) The spatial fluorescence distribution in steady state is measured. Three curves with 5 nM, 80 nM and 300 nM are shown exemplarily in FIG. 25A. The signal changes dramatically from fluorescence decrease to a fluorescence increase. Integration of the fluorescence profile up to 80 μm (distance from the heated centre) allows to determine the number of complexes in solution, as shown in FIG. 25B. The concentration of free Interleukin 4 can be calculated plotted versus the concentration formed complex. These data can be fitted to determine the K_(D).

FIG. 26 shows a single molecule binding to nanoparticles. The Soret coefficient of nanocrystals in complex with PEG molecules is measured by evaluating the concentration profile in steady state. The Soret coefficient increases linearly with the number of PEG molecules covalently bound to the particle. PEG molecules with a higher molecular weight show a steeper increase in the Soret coefficient. The PEG molecules shown here are comparable in size to proteins or short DNA molecules, which can be detected in the same manner.

FIG. 27 shows a further embodiment of a device according to the present invention. The receiving means for receiving the sample probe 50 is a capillary 40 with inner diameter 5 μm to 500 μm such that the thickness of the sample probe is small in the direction perpendicular to the laser beam. The first valve 41 and the second valve 42 are provided for the controlled input/output of the sample probe 50 in/from the capillary 40. The capillary is mounted on a solid support 43, preferably a material with good thermal conductivity, e.g. aluminium, copper. The Peltier-element 44 is mounted on the solid support 43 such that the capillary 40 can be cooled.

FIG. 28A-B shows the characterization of protein conformation. Thermo-optical characterization provides the means to characterize the conformation of a protein in solution. In FIG. 28A the temperature of a solution containing Bovine Serum Albumin (BSA) is cooled down to 0° C. Starting from this temperature the Soret coefficient is measured at different temperatures which are increased in a stepwise manner up to 60° C. The Soret coefficient is negative, up to values close to thermal denaturation, where a sudden jump to positive Soret coefficients is observed. At physiological temperatures (30-40° C.) the Soret coefficient does not change much. In this temperature range the protein has to have similar properties to perform its tasks. Because of the tight relation between structure and function, the conformation is preserved in this temperature range. The results are confirmed by the experiment shown on the right (FIG. 28B), which starts at high temperatures. The Soret Coefficients are still positive below 50° C. since the measurements are faster than the time the protein needs for refolding. After a certain time span (i.e. 20 minutes) the values reach the negative Soret coefficients obtained in the measurements started at low temperatures. A following temperature increase reproduces the negative Soret coefficient measured in the experiment starting at low temperatures.

FIG. 29 shows measurements with a sample of fluorescently labelled bovine serum albumin (BSA). A sample of fluorescently labelled bovine serum albumin (BSA) has been split in two parts. One is only exposed to ambient temperatures, while the other half is heated up to 100° C. for several minutes. The thermo-optical properties of both samples (native and denatured) are measured at different power of the infrared laser (i.e. maximal temperature increase of 5° or 10° C.). As can be seen from the figure, the fluorescence of the denatured protein is lower than the fluorescence of the native protein. This is explained as follows. The fluorescence dye of both samples shows the same decrease in fluorescence due to the increase in temperature (i.e. temperature sensitivity of the fluorescence). But the denatured protein shows a positive thermophoretic mobility (i.e. moves to the cold), while the native protein has a negative thermophoretic mobility (i.e. moves to the hot). The accumulation at elevated temperatures is the reason, why the decrease in fluorescence is lower for the native protein, while the denatured protein is, in addition to temperature dependency, depleted from the region of elevated temperature. The differences between both samples is further increases by raising the temperature (i.e. maximum temperature of 10° C.), since positive and negative thermophoresis is enhanced.

FIG. 30 measurements with a sample of fluorescently labelled bovine serum albumin (BSA). A sample of fluorescently labelled bovine serum albumin (BSA) has been split in two parts. One is only exposed to ambient temperatures, while the other half is heated up to 100° C. for several minutes (i.e. irreversibly denatured). The thermo-optical properties of both samples (native and denatured) are measured at 800 mA power of the infrared laser (i.e. maximal temperature increase of 20° C.). As can be seen from the figure, the fluorescence of the denatured protein is lower than the fluorescence of the native protein. This is explained as follows. The fluorescence dye of both samples shows the same decrease in fluorescence due to the increase in temperature (i.e. temperature sensitivity of the fluorescence). But the denatured protein shows a positive thermophoretic mobility (i.e. moves to the cold), while the native protein has a negative thermophoretic mobility (i.e. moves to the hot). The accumulation at elevated temperatures is the reason, why the decrease in fluorescence is lower for the native protein, while the denatured protein is, in addition to temperature dependency, depleted from the region of elevated temperature. Interestingly, by approaching the denaturing temperature (i.e. 50° C.) of the protein the amplitudes of the native and denatured protein approach each other an are essentially the same. This means that by measuring the amplitude of the fluorescence change an comparison to the reference sample allows to detect the melting temperature of a protein and to discriminate between native and denatured form of a protein. And to detect a shift in melting temperature due to interactions of the protein with other biomolecules or small molecules (e.g. drug candidates).

FIG. 31A-B shows the measurement of the thermo-optical properties of two samples with/of Green Fluorescent protein (GFP). The thermo-optical properties of two samples of Green Fluorescent protein (GFP) are measured. In the first sample (FIG. 31A) only GFP is present, while in the second sample (FIG. 31B) the GFP is mixed with a 2 fold excess of an antibody fragment, specifically binding to GFP. In both cases, first the fluorescence is recorded without laser heating. Then the fluorescence excitation is turned off and the IR-laser radiation is turned on. The laser is turned off after a few seconds of heating and the fluorescence excitation is turned on at the same time. The relaxation of the spatial fluorescence distribution (i.e. concentration distribution) to a homogeneous state is recorded for a few seconds. As can be observed from the figure, in the sample with the two interacting species (i.e. GFP and the antibody fragment) the fluorescence profile needs more time to relax. This is explained by slower diffusion of the larger complex. The time evolution of the fluorescence profile is analyzed via a software tool to determine the diffusion constant. By using the Stokes-Einstein relation, an hydrodynamic radius is attributed to the diffusion constant. In case of the free GFP this is 5 nm and the complex has an radius of 10 nm.

FIG. 32A-C shows a measurement of a particle which is trapped in a potential well created by a spatial temperature distribution. (A) A particle is trapped in a potential well created by a spatial temperature distribution. For silica particles the well is deepest at high temperatures. The fluctuations are recorded via a CCD camera (at t=1 s, 2 s, 3 s, 4 s, 5 s, 6 s, 7 s) and (B) the positions are tracked by Software with nanometer resolution. (C) A histogram is calculated from the positional information. The width of the distribution is very sensitive to the thermo-optical properties of the particle. If molecules bind to the surface of the particle, the effective potential for the bead changes and the amplitude of the fluctuations increases or decreases. By observing the amplitude change over time, a kinetic binding curve can be measured.

FIG. 33A-D shows a time series of the thermophoretic motion of silica beads in a microfluidic chamber. Time series of the thermophoretic motion of silica beads in a microfluidic chamber. In the beginning (image A), without laser heating, the beads are almost equally distributed. The black circle shows the position of the laser focus. The following images (B-D) show the development of the particle distribution in the next three seconds after the heating laser is turned on. The particles are attracted by the heat source and accumulate at the point of highest temperature. The accumulation if observed because these particles have a negative thermophoretic mobility. A Particle with positive thermophoretic mobility can be trapped by heating e.g. a circle around it.

FIG. 34 shows another example for the “Optothermal Trap”. Another Example for the “Optothermal Trap”: Several 1 μm beads are trapped at the bright spot at the centre of the image. The chamber is moved whereas the laser focus was kept fixed. The image is an addition of about 30 single images. As one can see, all beads were moved with the chamber, so the addition of the single images results in lines for the single particles. The trapped beads were hold on one position. No movement of the trapped beads was detected. The halo and the high intensity of the trapped beads is the result of the addition of the single images.

FIG. 35A shows the determination of the Soret Coefficient of complexes of nanocrytals (=quantum dot QD) and biomolecules. The Soret Coefficient of complexes of nanocrytals (=quantum dot QD) and biomolecules is determined by relating the spatial concentration distribution to a spatial temperature distribution. Three different samples have been analyzed. First a nanocrystal without protein modification is measured (QD), followed by a sample nanocrystals modified with the protein streptavidin (QD+Strep.)(approximately 5 proteins per nanocrystal). By binding the protein to the nanocrystal, the Soret-Coefficient is strongly increased. By adding a single stranded DNA to the sample (one DNA per Particle), the Soret coefficient is increased further (QD+Strep.+DNA). FIGS. 35B-D show the configurations of the complexes.

FIG. 36 shows a further embodiment of the invention according to FIG. 20, wherein a stage 43 carrying a temperature control element 44 and the chamber 45 is connected to the optical system via connectors 48. The optical system 1 may also be also comprise a TIRF (total internal reflection fluorescence) objective so that Thermophoresis can be measured in direction of the Laser beam.

FIG. 37 shows a further embodiment of the invention according to FIG. 20, wherein a shutter 33 and a line forming module 12, preferably a lens system comprising one, two or more lenses or more preferably a Powell lens are added to control the IR laser radiation and wherein the emission filter 7 is replaced by an optical instrument 22 which may be a spectrograph, polychromator or monochromator or combinations of one or more of these, e.g. an optical instrument which transforms different intervals of wavelength/frequency of light into different intervals of angles/distances or different places for example on a CCD.

FIG. 38 shows an example for a lipid bilayer model system. A fraction of the layer constituting lipids is coupled to a surface (e.g. via a sulfhydryl-peptide to a gold surface) Transmembrane proteins or membrane associated proteins are inserted into the lipid bilayer. In addition also soluble proteins may be present in the aqueous solution on top of the membrane. By infrared laser absorption of the aqueous solution a temperature gradient can be established within the membrane. This way the thermo-optical properties like stability, interaction and conformation may can be measured for a fluorescently labelled compound (i.e. lipid, membrane protein or soluble protein).

FIG. 39A-B shows thermophoresis and thermophoretic trapping of lipid vesicles. The images (200×200 μm) show a solution of unilamelar vesicles, without (A) and after 10 seconds of infrared laser heating. (A) shows a uniform distribution of the vesicles. The infrared laser heats the solution locally to a maximum temperature of 15° C. above room temperature of 20° C. As (B) shows the local temperature increase attracts the vesicles (i.e. negative thermophoresis) and confines their position to a region close to the center of the heat spot. The region around the heat spot is depleted of vesicles. Vesicles closer to the edge of the field of view experience only a small gradient an are not attracted within the time span of 10 seconds. Broadening of the temperature profile would also attracted these particles much faster.

EXAMPLES

The following detailed example illustrates the invention without being limiting.

Example 1: Thermodiffusion

Molecules drift along temperature gradients, an effect called thermophoresis, Soret-effect or thermodiffusion. In liquids, its theoretical foundation is subject of a long standing debate. Using a new all-optical microfluidic fluorescence method, we present experimental results for DNA and polystyrene beads over a large range of particle size, salt concentration and temperature. The data supports a unifying theory based on the solvation entropy. Stated in simple terms, the Soret coefficient is given by the negative solvation entropy, divided by kT. The theory predicts the thermodiffusion of polystyrene beads and DNA without any free parameters. We assume a local thermodynamic equilibrium of the solvent molecules around the molecule. This assumption is fulfilled for moderate temperature gradients below the fluctuation criterion. Above this criterion, thermodiffusion becomes non-linear. For both DNA and polystyrene beads, thermophoretic motion changes sign at lower temperatures. This thermophilicity towards lower temperatures is attributed to an increasing positive entropy of hydration, whereas the generally dominating thermophobicity is explained by the negative entropy of ionic shielding. The understanding of thermodiffusion sets the stage for detailed probing of solvation properties of colloids and biomolecules. For example, we successfully determine the effective charge of DNA and beads over a size range which is not accessible with electrophoresis.

Introduction.

Thermodiffusion has been known for a long time, but its theoretical explanation for molecules in liquids is still under debate. The search for the theoretical understanding is motivated by the fact that thermodiffusion in water might lead to powerful all-optical screening methods for biomolecules and colloids. Equally well, thermodiffusion handles and moves molecules all-optically and therefore can complement well established methods as for example electrophoresis or optical-tweezers. For the latter, forces of optical tweezers scale with particle volume and limit this method to particles only larger than 500 nm. Electrophoresis does not suffer from force limitations, but is difficult to miniaturize due to electrochemical reactions at the electrodes.

On the other hand, thermodiffusion allows the microscale manipulation of even small particles and molecules. For example, 1000 bp DNA can be patterned arbitrarily in bulk water (FIG. 7). The temperature pattern “DNA”, heated by 2 K, was written into a water film with an infrared laser scanning microscope. The concentration of 1000 bp DNA was imaged using a fluorescent DNA tag. In an overall cooled chamber at 3° C., DNA accumulates towards the heated letters “DNA” (negative Soret effect) whereas at room temperature DNA is thermophobic (positive Soret effect) as seen by the dark letters.

In the past, the apparent complexity of thermodiffusion prevented a full theoretical description. As seen for DNA in FIG. 13, molecules characteristically deplete from regions with an increased temperature, but they can also show the inverted effect and accumulate³. Moreover, the size scaling of thermodiffusion recorded by thermal field flow fractionation (ThFFF) showed fractional power laws with a variety of exponents which are hard to interpret⁴. The latter effect was resolved recently by revealing nonlinear thermophoretic drift for the strong thermal gradients used in ThFFF.

A variety of methods were used to measure thermodiffusion, mostly in the nonaqueous regime. They range from beam deflection^(3,7), holographic scattering^(8,9), electrical heating to thermal lensing. Recently we have developed a fluorescence microfluidic imaging technique^(13,14) which allows the measurement of thermodiffusion over a wide molecule size range without artifacts induced by thermal convection. Highly diluted suspensions can be measured and therefore particle-particle interactions do not have an influence. We only apply moderate temperature gradients. In the following we used this method to confirm a straightforward theoretical explanation of thermodiffusion.

Theoretical Approach.

For diluted concentrations, it is generally assumed that the thermodiffusive drift velocity {right arrow over (v)} depends linearly on the temperature gradient ∇T with a proportionality constant which equals the thermodiffusion coefficient D_(T): {right arrow over (v)}=−D_(T)∇T. In steady state, thermodiffusion is balanced by ordinary diffusion. Constant diffusion and thermodiffusion coefficients both lead to an exponential depletion law¹⁶ c/c₀=exp[−(D_(T)/D)(T−T₀)], with the concentration c depending on the temperature difference T−T₀ only. The concentration c is normalized by the boundary condition of the concentration c₀ with temperature T₀. The Soret coefficient is defined as ratio S_(T)=D_(T)/D which determines the magnitude of thermodiffusion in the steady state. While the above exponential distribution could motivate an approach based on Boltzmann equilibrium statistics, it is commonly argued that thermodiffusion without exception is a local non-equilibrium effect that requires fluid dynamics, force fields or particle-solvent potentials¹⁷⁻²⁰. However, in two previous papers¹⁶ we demonstrated that for moderate temperature gradients, the thermal fluctuations of the particle are the basis for a local equilibrium. This allows the description of the thermodiffusive steady state by a succession of local Boltzmann laws, yielding c/c₀=exp[(G(T₀)−G(T))/kT] with G the Gibbs-free enthalpy of the single particle-solvent system. Such an approach is only valid if the temperature gradient ∇T is below a threshold ∇T<(aS_(T))⁻¹ which is given by the particle fluctuations with the hydrodynamic radius a and Soret coefficient S_(T), as shown recently. For larger temperature gradients, thermodiffusive drift is nonlinearly dependent on the temperature gradient. In the present study, temperature gradients below this limit were used so that thermodiffusion is measured at local thermodynamic equilibrium conditions.

Local thermodynamic equilibrium allows the derivation of a thermodynamic foundation of the Soret coefficient. The local Boltzmann distribution relates small concentration changes δc with small Gibbs-Free Energy differences: δc/c=−δG/kT. We equate this relation with a locally, linearized thermodiffusion steady state given by δc/c=−S_(T)δT and thus find the Soret coefficient by the temperature derivative of G S _(T) =D _(T) /D=(kT)⁻¹ ×∂G/∂T  (1)

Whereas above relation is sufficient for the following derivation, it can be generalized by locally applying the thermodynamic relation dG=−SdT+Vdp+μdN. For single particles at a constant pressure we find that the Soret coefficient equals the negative entropy of the particle-solvent system S according to S_(T)=ΔS/kT. This relation is not surprising since the entropy is by definition related with the temperature derivative of the free enthalpy.

The above general energetic treatment is inherent in previously described approaches based on local equilibrium²², including the successful interpretation of thermoelectric voltages of diluted electrolytes^(24,25) which are described by energies of transfer. Recently, the non-equilibrium approach by Ruckenstein was applied to colloids²⁷ with the characteristic length l assigned to the Debye length λ_(DH). If instead it would assigned the characteristic length according to l=2a/3 with the particle radius a, the Ruckenstein approach would actually confirm the above local equilibrium relation (1) for the Soret coefficient. Measurements on SDS micelles²⁷ appeared to confirm this non-equilibrium approach, but for the chosen particles the competing parameter choices l=2a/3 and l=λ_(DH) yielded comparable values. Thus the experiments could not distinguish between the competing theories.

Here, we will use the above local equilibrium relations to derive the Soret coefficient for particles larger than the Debye length in aqueous solutions and put the results to rigorous experimental tests. Two contributions dominate the particle entropy S in water (FIG. 8a ): the entropy of ionic shielding and the temperature sensitive entropy of water hydration. The contribution from the entropy of ionic shielding is calculated with the temperature derivative of the Gibbs-free enthalpy^(27,28) G_(ionic)=Q_(eff) ²λ_(DH)/[2Aεε₀] with the effective charge Q_(eff) and particle surface A. Alternatively, this enthalpy can be interpreted as an electrical field energy G_(ionic)=Q_(eff) ²/[2C] in the ionic shielding capacitor C. We neglect the particle-particle interactions since the fluorescence approach allows the measurement of highly diluted systems. To obtain the Soret coefficient, temperature derivatives consider the Debye length λ_(DH)(T)=√{square root over (ε(T)ε₀kT/(2e²c_(S)))} and the dielectric constant ε(T). Both temperature derivatives give rise to a factor β=1−(T/ε)∂ε/∂T. The effective charge Q_(eff) is largely temperature insensitive which was confirmed by electrophoresis independently²⁹. Such a dependence would be unexpected as the strongly adsorbed ions dominate the value of the effective charge. Experimentally, we deal with colloids exhibiting flat surfaces, i.e. the particle radius is larger than λ_(DH). In this case charge renormalization does not play a role and we can introduce an effective surface charge density σ_(eff)=Q_(eff)/A per molecule area A. From the temperature derivation according to equation (1), the ionic contribution to the Soret coefficient is S_(T) ^((ionic))=(Aβσ_(eff) ²λ_(DH))/4εε₀kT²). A similar relation was derived for charged micelles recently²³, however without considering the temperature dependence of the dielectric coefficient ε. Next, the contribution to the Soret coefficient from the hydration entropy of water can be directly inferred from the particle area specific hydration entropy s_(hyd)=S_(hyd)/A, namely S_(T) ^((hyd))=−As_(hyd)(T)/kT. Finally, the contribution from the Brownian motion is derived as S_(T)=1/T by inserting the kinetic energy of the particle G=kT into equation (1). However this contribution is very small (S_(T)=0.0034/K) and can be neglected for the molecules under consideration. The contributions from ionic shielding and hydration entropy add up to:

$\begin{matrix} {S_{T} = {\frac{A}{kT}\left( {{- s_{hyd}} + {\frac{{\beta\sigma}_{eff}^{2}}{4{ɛɛ}_{0}T} \times \lambda_{DH}}} \right)}} & (2) \end{matrix}$

The Soret coefficient S_(T) scales linearly with particle surface A and Debye length λ_(DH). We test equation (2) by measuring S_(T) versus the salt concentration, temperature and molecule size. In all cases thermodiffusion is quantitatively predicted without any free parameters. We used fluorescence single particle tracking to follow carboxyl modified polystyrene (PS) beads (Molecular Probes F-8888) of 1.1 μm and 0.5 μm diameter at 25 attomolar concentration, dialyzed into 0.5 mM Tris-HCl at pH 7.6. Thermodiffusion of particles ≤0.2 μm is measured by the fluorescence decrease that reflects the bulk depletion of the particles¹³. The chamber thickness of 20 μm damped the thermal convection to negligible speeds¹⁶. The experimental design also excludes thermal lensing and optical trapping¹⁶. Debye lengths λ_(DH) were titrated with KCl (see supplementary materials).

Salt Dependence.

FIG. 8b shows the Soret coefficients of polystyrene beads with different sizes versus λ_(DH). The Soret coefficients scale linearly with a small intercept at λ_(DH)=0 and confirm the λ_(DH)-dependence of equation (2). For smaller diameters of the beads the Soret coefficients scale with the particle surface area A (FIG. 8) as expected from equation (2). To check whether equation (2) also quantitatively explains the measured Soret coefficients, we inferred the effective charge of the beads by electrophoresis (see supplementary material). Using 40 nm beads with identical carboxyl surface modifications at λ_(DH)=9.6 nm, we fluorescently observed free-flow electrophoresis and corrected for electroosmosis, finding an effective surface charge density of σ_(eff)=4500±2000e/μm². This value is virtually independent from the used salt concentrations²⁹. Using this inferred effective charge, equation (2) fits the Soret coefficient for various bead sizes and salt concentrations well (FIG. 8b , solid lines).

The intercept S_(T)(λ_(DH)=0), where ionic contributions are zero, also scales with particle surface and is described by a hydration entropy per particle surface of s_(hyd)=1400 J/(molKμm²). The value matches the literature values for similar surfaces reasonably well^(30,31). For example, Dansyl-Alanine, a molecule with surface groups comparable with polystyrene beads, was measured to have a hydration entropy³⁰ of −0.13 J/(molK) at a comparable temperature. Linear scaling with its surface area by using a radius of a=2 nm results in a value of S_(hyd)=−2500 J/(molKμm²), in qualitative agreement with our result. The hydration entropy is a highly informative molecule parameter which is notoriously difficult to measure, yielding an interesting application for thermodiffusion.

Temperature Dependence.

Hydration entropies S_(hyd) in water are known to increase linearly with decreasing temperature³⁰⁻³². Since the slope of the ionic contribution of S_(T) versus λ_(DH) is with high precision temperature insensitive for water (β(T)/(εT²)≅const), only the intercept is expected to decrease as the overall temperature of the chamber is reduced. This is indeed the case, as seen from the temperature dependence of beads with 1.1 μm diameter (FIG. 9a , T=6 . . . 29° C.). We infer from the intercept S_(T)(λ_(DH)=0) that the hydration entropy changes sign at about 20° C. As seen for DNA in FIG. 7, hydration entropy can dominate thermodiffusion at low temperatures and move molecules to the hot (D_(T)<0).

The properties of hydration entropy lead to a linear increase of S_(T) over temperatures at fixed salt concentration as measured for 1.1 μm beads and DNA (FIG. 9b ). We normalize S_(T) by dividing with S_(T)(30° C.) to compensate for molecule surface area. The slopes of S_(T) over temperature differ between beads and DNA. However the slope does not differ between DNA of different size (50 base pairs versus 10000 base pairs). Based on equation (2), this is to be expected since the temperature dependence of the hydration entropy only depends on the type of surface of the molecule, not its size. We measured the diffusion coefficients of the DNA species at the respective temperature independently. Within experimental error, changes in the diffusion coefficient D match with the change of the water viscosity without the need to assume conformational changes of DNA over the temperature range. Please note that the change of the sign of the DNA Soret coefficient is situated near the point of maximal water density only by chance. There, the two entropic contributions balance. For polystyrene beads at λ_(DH)=2 nm for example, the sign change is observed at 15° C. (FIG. 9a ). An increased Soret coefficient over temperature was reported for aqueous solutions before³, however with a distinct nonlinearity which we attribute to remnant particle-particle interactions.

Size Dependence of the Beads.

The Soret coefficient was measured for carboxyl modified polystyrene beads in diameter ranging from 20 nm to 2 μm (Molecular Probes, F-8888, F-8795, F-8823, F-8827). Beads of diameter 0.2 μm, 0.1 μm, 0.04 μm and 0.02 μm were diluted to concentrations of 10 pM, 15 pM, 250 pM and 2 nM, and its bulk fluorescence was imaged over time to derive D_(T) and D^(13,16) from the depletion and subsequent back-diffusion. Larger beads with a diameter of 1.9 μm, 1.1 μm and 0.5 μm were diluted to concentrations of 3.3 aM, 25 aM and 0.2 pM, and measured with single particle tracking⁶. The solutions were buffered in 1 mM Tris pH 7.6 with λ_(DH)=9.6 nm. In all cases interactions between particles can be excluded. Care was taken to keep the temperature gradient in the local equilibrium regime.

We find that the Soret coefficient scales with particle surface over four orders of magnitude (FIG. 10a ). The data is described well with equation (2) with an effective surface charge density of σ_(eff)=4500 e/μm² and neglected hydration entropy contribution. The 5-fold too low prediction for the smallest particle (20 nm diameter) can be explained by charge renormalization since its radius is smaller than λ_(DH).

The diffusion coefficient D for spheres is given by the Einstein relation and scales inversely with radius D∝1/a. Inserting equation (2) into S_(T)=D_(T)/D, the thermodiffusion coefficient D_(T) is expected to scale with particle radius a. This is experimentally confirmed over two orders of magnitude (FIG. 10b ). These findings of ours contradict several theoretical studies claiming that D_(T) should be independent of particle size^(17-20,27), based on ambiguous experimental results from thermal field flow fractionation (ThFFF)⁴ which were probably biased by nonlinear thermodiffusion in large thermal gradients.

Size Dependence of DNA.

Whereas polystyrene beads share a very narrow size distribution as a common feature with DNA molecules, beads are a much less complicated model system. Beads are rigid spheres which interact with the solvent only at its surface. In addition, the charges reside on the surface where the screening takes place. Thus the finding that thermodiffusion of flexible and homogeneously charged DNA is described equally well described with equation (2) is not readily expected and quite interesting (FIG. 10c,d ).

We measured DNA sizes with 50 base pairs to 48502 base pairs in 1 mM TRIS buffer (λ_(DH)=9.6 nm) at low molecule concentrations between 1 μM (50 bp) and 1 nM (48502 bp). Only every 50th base pair was stained with the TOTO-1 fluorescent dye. The diffusion coefficient was measured by back-diffusion after the laser was turned off and depends on the length L of the DNA in a non-trivial way. The data is well fitted with a hydrodynamic radius scaling a ∝L^(0.75). This scaling represents an effective average over two DNA length regimes. For DNA molecules longer than approximately 1000 bp, a scaling of 0.6 is found³³ whereas shorter DNA scales with an exponent of ≈1 (see supplementary material).

We can describe the measured Soret coefficient over three orders of magnitude of DNA lengths with equation (2) if we assume that effective charge of the DNA is shielded at the surface of a sphere with the hydrodynamic radius a. Due to the low salt concentration (λ_(DH)=9.6 nm), such globular shielding is reasonable. Not only is the experimentally observed scaling of the Soret coefficient with the square root of its length correctly predicted based on equation (2) (S_(T)∝Q_(eff) ²/a²∝L²/L^(1.5)∝L^(0.5)), also the Soret coefficient is fully described in a quantitative manner (FIG. 10c , solid line), with an effective charge of 0.12 e per base, matching well with literature values³⁴ ranging from 0.05 e/bp to 0.3 e/bp.

As shown in FIG. 10d , the thermodiffusion coefficient for DNA drops with DNA length according to D_(T)=DS_(T)∝Q_(eff) ²/a³∝L²/L^(2.25)∝L^(−0.25). Thus, shorter DNA actually drifts faster in a temperature gradient than longer DNA. It is important to point out that this finding is in no contradiction to experimental findings of a constant D_(T) over polymer length in non-aqueous settings⁹. According to equation (1), the thermodynamic relevant parameter is the Soret coefficient, determined by the solvation energetics. The argument²⁰ that polymers have to decouple into monomers to show a constant D_(T) merely becomes the special case where the solvation energetics determine both S_(T) and D with equal but inverted size scaling. In accordance with our local energetic equilibrium argument, S_(T) and not D_(T) dominates thermodiffusion also for non-aqueous polymers near a glass transition³⁵. Here, S_(T) is constant whereas D_(T) and D scale according to an increased friction. However for a system of DNA in solution, where long ranging shielding couples the monomers, a constant D_(T) over polymer length cannot be assumed a priori (FIG. 10d ).

Effective Charge.

The effective charge Q_(eff) is a highly relevant parameter for colloid science, biology and biotechnology. So far it only could be inferred from electrophoresis, restricted to particles smaller than the Debye length (a≤3λ_(DH))³⁶. Unfortunately, many colloids are outside this regime. As shown before, a similar size restriction does not hold for thermodiffusion. In many cases, the hydration entropy s_(hyd) contributes less than 15% (FIG. 8) and can be neglected at moderate salt levels. Thus we can invert equation (2) to obtain the effective charge Q_(eff) for spherical molecules from

$\begin{matrix} {Q_{eff} = {\frac{2T^{2}}{3\eta\; D}\sqrt{\frac{{ɛɛ}_{0}k^{3}S_{T}}{{\beta\pi\lambda}_{DH}}}}} & (3) \end{matrix}$

The effective charge derived from thermodiffusion measurements of polystyrene beads and DNA is plotted in FIG. 11 over several orders of magnitude in size. The effective charge of beads scales linearly with particle surface with a slope confirming the effective surface charge density of σ_(eff)=4500 e/μm² which was inferred from electrophoresis only for small particles. Average deviations from linear scaling are below 8% (FIG. 11a ). The effective charge inferred from thermodiffusion measurements of DNA using equation (3) scales linearly with DNA length with an effective charge of 0.12 e per base pair. The length scaling is confirmed over four orders of magnitude with an average error of 12% (FIG. 11b ). Thus thermodiffusion can be used to infer the effective charge with low errors for a wide range of particle sizes. This is even more interesting for biomolecule characterization since measurements of thermodiffusion can be performed all-optically in picoliter volumes.

Conclusion

We describe thermodiffusion, the molecule drift along temperature gradients, in liquids with a general, microscopic theory. Applied to aqueous solutions, this theory predicts thermodiffusion of DNA and polystyrene beads with an average accuracy of 20%. We experimentally validate major parameter dependencies of the theory: linearity against screening length λ_(DH) and molecule hydrodynamic area A, quadratic dependence on effective charge and linearity against temperature. Measurements of thermodiffusion can be miniaturized to micron scale with the used all-optical fluorescence technique and permits microscopic temperature differences to manipulate molecules based on their surface properties (FIG. 7). The theoretical description allows to extract the solvation entropy and the effective charge of molecules and particles over a wide size range.

Infrared Temperature Control.

The temperature gradients used to induce thermodiffusive motions were created by aqueous absorption of an infrared laser at 1480 nm wavelength and 25 mW power (Furukawa). Water strongly absorbs at this wavelength with an attenuation length of κ=320 μm. The laser beam was moderately focussed with a lens of 8 mm focal distance. Typically, the temperature in the solution was raised by 1-2 K in the beam center with a 1/e² diameter of 25 μm, measured with the temperature dependent fluorescence signal of the dye BCECF¹³. Thin chamber heights of 10 μm to 20 μm and moderate focussing removed possible artifacts from optical trapping, thermal lensing and thermal convection¹³. For temperature dependent measurements both the objective and the microfluidic chip were tempered with a thermal bath. Imaging was provided from an AxioTech Vario fluorescence microscope (Zeiss), illuminated with a high power LED (Luxeon) and recorded with the CCD Camera SensiCam QE (PCO).

Molecules.

Highly monodisperse and protein-free DNA of 50 bp, 100 bp, 1000 bp, 4000 bp, 10000 bp and 48502 bp (Fast Ruler, fragments and λ-DNA, Fermentas) were diluted to 50 μM base pair concentration, i.e. the molecule concentration was between 1 μM (50 bp) and 1 nM (48502 bp). DNA was fluorescently labeled by the intercalating TOTO-1 fluorescent dye (Molecular Probes, Oregon) with a low dye/base-pair ratio of 1/50. Carboxyl modified polystyrene beads of diameter 2 μm, 1 μm, 0.5 μm, 0.2 μm, 0.1 μm, 0.04 μm and 0.02 μm (F-8888, F-8823, F-8827, F-8888, F-8795, F-8823, F-8827, Molecular Probes) were dialyzed (Eluta Tube mini, Fermentas) in aq. dest. and diluted in 1 mM Tris pH 7.6 to concentrations between 3.3 aM (2 μm) and 2 nM (0.02 μm).

Concentration Imaging Over Time.

Either the method of concentration imaging¹³ or single particle tracking were used to measure thermodiffusion at low concentrations, namely below 0.03 g/1 for DNA and 10⁻⁵ g/l for beads. At higher concentrations, we found profound changes of thermodiffusion coefficients. DNA and polystyrene beads smaller than 0.5 μm diameter concentration were imaged over time¹³ by bright field fluorescence with a 40× oil immersion objective. Concentrations inferred after correcting for bleaching, inhomogeneous illumination and temperature dependent fluorescence¹³ were fitted with a finite element theory. The model captures all details of both thermodiffusive depletion and backdiffusion to measure D_(T) and D independently (see supplementary material). Measurements were performed in microfluidic chips 10 μm in height with PDMS on both sides¹³.

Single Particle Tracking.

Polystyrene particles larger than 0.5 μm in diameter were measured by single particle tracking due to the slow equilibration time and risk that steady state depletion is disturbed by thermal convection. The thermodiffusive drift was imaged with a 32× air objective at 4 Hz at an initial stage of depletion in a 20 μm thick chamber. Averaging over the z-position of the particles removed effects from thermal convection. The drift velocity versus temperature gradient of 400 tracks were linearly fitted by v=−D_(T)∇T to infer D_(T). The diffusion coefficients D of the particles were evaluated based on their squared displacement, matching within 10% the Einstein relationship.

The present invention is not limited by what has been particularly shown and described hereinabove. Rather the scope of the present invention includes both combinations and sub-combinations of the features described hereinabove as well as modifications and variations thereof which would occur to a person skilled in the art upon reading the foregoing description and which are not in the prior art.

Infrared Heating.

The temperature gradients used to induce thermophoretic motions are created by aqueous absorption of a fiber coupled infrared solid state laser (Furukawa FOL1405-RTV-317), with a wavelength of 1480 nm and a maximum power of 320 mW typically used at 25 mW. Water strongly absorbs at this wavelength with an attenuation length of κ=320 μm. The infrared light is coupled out of the fiber to form a parallel beam with an 1/e² diameter of 1 mm. The beam position in the x/y plane can be adjusted by two galvanoometrically controlled infrared mirrors (Cambridge Technology 6200-XY Scanner with Driver 67120). The laser beam is focussed from below the object stage by an infrared corrected aspheric lens with 8 mm focal distance (Thorlabs, C240TM-C). Typically, temperature was increased by only 2 K in the heated focus.

Temperature Measurement.

The temperature gradient was measured via the temperature dependent fluorescence signal of the dye BCECF, diluted to 50 μM in 10 mM TRIS buffer. Details of bleaching correction and temperature extraction were described previouslyl¹³. From the total temperature dependence of BCECF of −2.8%/K, only −1.3%/K stems from pH drift of the used TRIS buffer. The remaining −1.5%/K are the result of thermodiffusion of the dye itself, measured to be S_(T)=0.015/K with the concentration over time method described below.

“DNA” Image.

Measurement of the “DNA” profile in FIG. 7 was performed in a 60 μm thick chamber between glass slides, imaged with a 10× objective and heated to 2 K along the letters “DNA” with laser scanning. The chamber was filled with a 50 nM solution of 1000 bp DNA stained with the intercalating fluorescent dye TOTO-1 (Molecular Probes) To switch from depletion to accumulation, the experiment was performed at room temperature or with the chamber cooled to 3° C., respectively.

Fluorescence Approaches.

Historically, methods used to measure thermodiffusion in liquids are based on changes in refractive index upon change in solute concentration²⁷. Inherently, this signal is small for low solute concentrations near the limit of non-interacting molecules, even for intricate detection methods like thermal lensing or holographic interference³⁸. Although operating at much smaller volume, the used fluorescent microfluidic approach¹³ allows concentrations of 0.03 g/l for DNA and reaches 10⁻⁵ g/l for single particle tracking. This is necessary, since for example in thermodiffusion of DNA, we see profound changes at higher concentrations.

Thermodiffusion from Concentration Over Time.

Both DNA and polystyrene beads smaller than 0.5 μm in diameter were measured by imaging molecule concentration over time by bright field fluorescence. A more basic steady state method was described previouslyl¹³. Here, we refined it with a numerics theory to infer diffusion coefficient D and Soret coefficient S_(T) independently.

Highly monodisperse and protein-free DNA of 50 bp, 100 bp, 1000 bp, 4000 bp, 10000 bp and 48502 bp (Fast Ruler, fragments and λ-DNA, Fermentas) were used for the length dependent measurements. The DNA was fluorescently labeled by the intercalating TOTO-1 fluorescent dye (Molecular Probes, Oregon) which shows 1000× fluorescence increase when bound to DNA. The dye/base-pair ratio was low (1/50) to avoid structural or charge artifacts from the bound dye. The fluorescence was observed with an 40× oil objective. DNA stock solutions were diluted to 50 μM base pair concentration, which corresponds to molecule concentrations between 1 μM (50 bp) and 1 nM (48502 bp), respectively. Polystyrene beads of diameter 0.2 μm, 0.1 μm, 0.04 μm and 0.02 μm (Molecular Probes, Oregon, F-8888, F-8795, F-8823, F-8827) were dialyzed (Fermentas, Eluta Tube mini) in aq. dest., and diluted in 1 mM Tris pH 7.6 to concentrations of 10 pM, 15 pM, 250 pM and 2 nM, respectively. DNA thermodiffusion measurements were performed in microfluidic chips 10 μm in height with PDMS on both sides¹³. They allow the measurement of small volumes in thin defined geometries. Polystyrene beads were sandwiched between a 1.25 mm thick polystyrene slide (Petri Dish, Roth) on the bottom and a plastic slide (170 μm thick, 1 cm×1 cm, Ibidi, Munich) on the top and sealed with nail polish. For temperature dependent measurements, both 40× oil objective and microfluidic chip were tempered from below with a thermal bath. Note that temperatures well below 0° C. can be achieved as the microfluidic geometry reduces the probability of water to freeze.

Concentration of DNA was inferred from fluorescence images, that were measured with a 40× oil objective (NA=1.3) on an AxioTech Vario fluorescence microscope (Zeiss), illuminated with a high power LED (Luxeon) and imaged with the CCD Camera SensiCam QE (PCO). The time series allows to correct for inhomogenous illumination and bleaching¹³. A time series dependent photobleaching means that an individual bleaching factor is determined for every image. This correction is of advantage for high precision measurements for protein thermophoresis. If single molecules were visible in fluorescence, time averaging was used to average over particle positions.

Radial profiles were taken over time and combined to a space-time image of both the thermophoretic depletion and back diffusion (FIG. 12a,c ). Fluorescence was adjusted for the temperature dependence of the TOTO-1 dye determined independently with a fluorescence spectrometer as −0.5%/K. Typically, temperature in the solution was raised by 1-2 K in the beam center with a 1/e² diameter of 25 μm.

The Soret coefficient S_(T) can be obtained from the steady state profile. Given the temperature at radius r obtained from temperature dependent fluorescence, the concentration c(r) can be fitted with the steady state thermophoretic profile¹³ given by c(r)=c ₀ e ^(−S) ^(T) ^((T(r)-T) ⁰ ⁾  (4)

with chamber temperature T₀ and bulk concentration c₀.

We can also obtain D and D_(T) independently by analyzing the build up and flattening of concentration profile over time after turning the infrared laser beam on or off, respectively. Theory was provided from finite element model in radial coordinates (FEMLab, Comsol) over time with boundary conditions of concentration obtained from the experiment. Comparison with experiment of the time course of thermophoretic depletion reveals D_(T) (FIG. 12c,d ) and from the time course after switching off the heating source the diffusion coefficient D is obtained (FIG. 12a,b ). Results for diffusion coefficients obtained for DNA molecules are shown in FIG. 13. Scaling of D for DNA larger than 1000 bp agree well with literature values and theoretical expectations³³. However for DNA molecules in the order of the persistence length (about 150 bp) the power law exponent of −0.6 does not precisely fit the measured values and a different scaling with an exponent of −1 is necessary. A good description of DNA diffusion coefficients in the size range analyzed throughout this work is achieved with an intermediate exponent of −0.75.

Screening length. Debye-Hüickel length was titrated by adding c_(S)=0 mM, 2 mM and 20 mM of KCl to c_(T)=1 mM of TRIS buffer at pH 7.6 and calculated from

$\begin{matrix} {\lambda_{DH} = \sqrt{\frac{{ɛɛ}_{0}{kT}}{2{e^{2}\left( {c_{S} + c_{T}} \right)}}}} & (5) \end{matrix}$

Changes in the effective charge of the molecules can be excluded at these monovalent salt concentrations. For largest values λ_(DH)=13.6 nm, solely 0.5 mM Tris-HCl pH 7.6 buffer was used.

Thermodiffusion Using Single Particle Tracking.

For fluorescent polystyrene particles of large size (2 μm, 1 μm, 0.5 μm, Molecular probes, Oregon, F-8888, F-8823 and F-8827) a different method had to be used due to the increasing visibility of single particles, slow equilibration time and the risk that steady state depletion is disturbed by thermal convection. Beads were dialyzed (Fermentas, Eluta Tube mini) in aq. dest., and diluted in 1 mM Tris pH 7.6 to concentrations of 3.3 aM, 25 aM and 0.2 pM, respectively. Thermodiffusion was measured in 20 μm thin chambers. A 1.25 mm thick polystyrene slide (Petri Dish, Roth, Karlsruhe) was chosen as for the bottom of the chamber, while a plastic slide (170 μm thick, 1 cm×1 cm, Ibidi, Munich) was taken as cover slip. The low thermal conductivity ensures constant temperature across the chamber. Chamber walls were made hydrophilic in a plasma cleaner (Harrick) for 10 min at 10 W electrical power. As a result, adsorption of polystyrene particles to the plastic is low even at high salt concentrations. Addition of 2 μl bead solution between the plastic sheets, followed by sealing the chamber with fast drying nail polish leads to reproducible chamber heights of 20 μm.

Imaging was provided with an AxioTech Vario fluorescence microscope (Zeiss), illuminated with a high power LED (Lirceon) and imaged with the CCD Camera SensiCamQE (PCO). The center of the plain of view was heated 8 K above room temperature with maximal temperature gradient of 0.2 K/μm and 1/e² spot radius of 50 μm Temperature was imaged with BCECF as described before in a separate chamber. In FIG. 14a finite element simulation of the experimental situation is shown using the commercial available Femlab software (Comsol). A 20 μm chamber is heated to 8 K in the center. Due to low heat conductivity of the PS walls the temperature profile is homogeneous throughout the chamber height (FIG. 14a ). Due to the thin chamber the convection speed is suppressed to negligible speeds of 5 nm/s at maximum (FIG. 14b ). The thermophoretic movement of the particles was imaged with a 32× air objective and recorded at 4 Hz. Particles all over the 20 μm chamber could be equally tracked with a custom-written Lab View program. Artefacts from toroidal thermal convection are averaged out to a high degree as convective attraction near the lower chamber wall cancelled with opposite convective repulsion near the top of the chamber. Typically the velocity of 400 tracks was plotted against radius and fitted with the drift velocity expected from thermodiffusion according to thermophoretic drift v=−D_(T)∇T to find the thermophoretic mobility D_(T). Thermal fluctuations of the tracks were evaluated based on their squared displacement to obtain the diffusion coefficient D of the particles, which matched within 10% the Einstein relationship D=kT/(6πηa). This is expected since even for the worst case, the chamber is 20-fold thicker than the diameter of the measured particles.

Electrophoresis.

The effective surface charge density was measured for 40 nm diameter beads by electrophoretic drift in 400 μm thin and 5 cm long chambers (Ibidi, Germany). The velocity profile throughout the chamber height at 5 V was taken from single particle tracking of 2 μm beads. At 80 μm height the electroosmotic flow in a tightly sealed chamber is zero³⁹. As expected the velocity of particles in this plane saturates for particles larger than 100 nm and is not related to effective charge³⁶. A high numerical aperture oil objective has been used to analyze the velocity of 40 nm particles at the chamber surface at the same conditions. The constant velocity difference between chamber surface and plane of zero electroosmotic flow measured before has been used to calculate the purely electrophoretic velocity.

Additional References referred to herein above in Example 1:

-   3. S. J. Jeon, M. E. Schimpf and A. Nyborg, Anal. Chem. 69,     3442-3450 (1997) -   4. P. M. Shiundu, G. Liu, and J. C. Giddings, Anal. Chem. 67,     2705-2713 (1995)6 -   7. B.-J. de Gans, R. Kita, B. Müller and S. Wiegand, J. Chem. Phys.     118, 8073 (2003). -   8. J. Rauch and W. Köhler, Phys. Rev. Lett. 88, 185901 (2002) -   9. S. Wiegand and W. Köhler, in Thermal Nonequilibrium Phenomena in     Fluid Mixtures, Springer, Berlin, 189 (2002) -   13. D. Braun and A. Libchaber, Physical Review Letters 89, 188103     (2002) -   14. S. R. de Groot, P. Mazur, Non Equilibrium Thermodynamics     (North-Holland, Amsterdam, 1969) -   16. A. H. Jr. Emery and H. G. Drickhammer, J. Chem. Phys. 23, 2252     (1955) -   17. J. S. Ham, J. Applied Physics, 31, 1853 (1960) -   18. K. I. Morozov, J. Experim. and Theor. Phys. 88, 944 (1999) -   19. M. E. Schimpf and S. N. Semenov, J. Phys. Chem. B 104, 9935     (2000) -   20. A. Voit, A. Krekhov, W. Enge, L. Kramer, and W. Köhler, Phys.     Rev. Lett. 92, 214501 (2005) -   22. S. Fayolle, T. Bickel, S. Le Boiteux and A. Wiirger, Phys. Rev.     Lett. 95, 208301 (2005) -   24. P. N. Snowdon, J. C. R. Turner, Trans. Faraday Soc. 56, 1409     (1960) -   25. E. Ruckenstein, J. Colloid Interface Sci. 83, 77 (1981) -   27. J. Israelachvili, Intermolecular & Surface Forces, 2nd edition,     Academic Press, 1992 -   28. W. Lin, P. Galletto and M. Borkovec, Langmuir 20, 7465-7473     (2004) -   29. D. Haidacher, A. Vailaya and C. Horváth, Proc. Natl. Acad. Sci.     93, 2290-2295 (1996) -   30. N. T. Southall, K. A. Dill and A. D. J. Haymet, J. Phys. Chem B     106, 521-533 (2002) -   31. B. Kronberg, M. Costas and R. Silveston, Pure & Applied     Chemistry 67, 897-902 (1995)

Example 2: Determination of Hydrodynamic Radius and Interaction Between Proteins

The thermo-optical characterization method of the present invention allows also to quantify the hydrodynamic radius of proteins and even more important of complexes of biomolecules which are not connected covalently to each other. Thermophoresis provides a comparably robust and precise way to measure the hydrodynamic radius of molecules from less than a nanometer up to a few microns. In comparison to the other thermo-optical characterization methods the precision of this method is not too sensitive on the geometry of measurement (e.g. height of the liquid layer) as it is the case for molecular interactions.

Data acquisition: A Typical Measurement can be Described as Follows:

Step 1:

A solution of fluorescently labeled molecules is brought into a microfluidic measurement chamber (e.g. capillary, microfluidic chip). Fluorescence is excited and recorded with spatial resolution for less than 5 seconds on a CCD device with a frame rate between 40 Hz down to 0.2 Hz (i.e. for fast diffusing molecules, a high frame rate is chosen). These image(s) provides the necessary information about the fluorescence level at 100% concentration. Then fluorescence excitation is turned off.

Step 2:

The Infrared laser heating is turned on. The immediately established local spatial temperature distribution causes the molecule drift to lower or higher temperatures, depending on the particular molecule to be analyzed. The laser is focused in a way that temperature gradients between 0.0 and 5 K/μm are achieved. The temperature gradient has been calibrated once and it is not necessary to repeat this calibration every time an experiment is performed. The maximal temperature elevation is below the temperature which is known to cause damage to the molecules or disintegrate their interaction. Depending on the thermophoretic properties of the molecules in the solution (i.e. if they move fast in a thermal gradient or slow) the infrared laser heats the solution for 5 seconds up to 100 seconds. After this period of time the infrared laser is turned off

Step 3:

After the spatial temperature distribution has vanished (typically 2-50 ms) the fluorescence excitation is turned on and fluorescence is recorded with the same frame rate used in the first step of fluorescence imaging. This time the redistribution of the molecules is imaged for 5 seconds up to 50 seconds. The exact time depends on the velocity with which the molecules diffuse (i.e. the time it takes them to equalize 90% of the concentration gradient established by thermophoresis).

Data Processing—Photobleaching:

The fluorescence images have to be corrected for photobleaching. Since there is no spatial temperature profile in the solution while fluorescence images are taken, the bleaching correction is possible with high precision (i.e. high precision is possible since the rate of photobleaching is temperature dependent).

Therefore, the fluorescence at a edge of the measurement chamber (i.e. a spot as far away from the heated center as possible), were thermophoresis during step 3 was negligible (i.e. for a person skilled in the art, this is where the temperature gradient during laser heating was lower than 0.001 K/μm), is evaluated to determine the photobleaching from the image series taken in step 3. If photobleaching is present, the fluorescence will decrease from image to image. The individual factor for each image is used to correct all images for bleaching. Another possibility is to calculate the bleaching for every single pixel from the images taken in Step 1. The bleaching rate per pixel can be used to correct every pixel from step 3 images for the photobleaching effect.

Data Processing—Inhomogeneous Illumination Correction and Normalization to 100% Concentration:

All images taken in step 4 are divided by a single or all images taken in step 2 and multiplied by 100. This way a correction for inhomogeneous illumination is achieved and the fluorescence is normalized to 100% concentration.

Data Processing—Determining the Hydrodynamic Radius:

From the first images of the step 4 image series the concentration distribution is extracted. A software tool evaluates the Diffusion coefficient (or multiple Diffusion coefficients in case of a mixture) which describes the experimentally measured relaxation of the concentration gradient. Using the Stokes-Einstein relation the hydrodynamic radius is inferred from the diffusion coefficient.

In particular, the above described experiment was conducted for a sample of GFP as follows:

The thermo-optical properties of two samples of Green Fluorescent protein (GFP) are measured with the devices of the present invention. 2 μl of GFP (5 μM, 1×PBS buffer) is pipetted on an object slide. The sample is sandwiched by pitting an cover slip (round 12 mm diameter) on top. The liquid spreads uniformly in between the glass surfaces an the chamber is sealed off by using nail polish. This prevents the liquid from rapid evaporation, which would in turn lead to a comparably strong flow of liquid in the chamber. This sample is palaced on a device shown in FIG. 1a . The measurement steps and data processing steps are performed as described above. The same experiment is performed with a second sample containing 5 μM GFP and 10 μM of a GFP binding antibody fragment, specifically binding to GFP. In both cases, first the fluorescence is recorded without laser heating. Then the fluorescence excitation is turned off and the IR-laser radiation is turned on (the maximum temperature is kept below 35° C. (i.e. about 15° C. above ambient temperature (about 20° C.) to avoid denaturation or damage to the protein). The laser is turned off after a few seconds of heating and the fluorescence excitation is turned on at the same time. The relaxation of the spatial fluorescence distribution (i.e. concentration distribution) to a homogeneous state is recorded for a few seconds. As can be observed from FIG. 31, in the sample with the two interacting species (i.e. GFP and the antibody fragment) the fluorescence profile needs more time to relax. This is explained by slower diffusion of the larger complex. The time evolution of the fluorescence profile is analyzed via a software tool (selfmade Software, Labview, National Instruments) to determine the diffusion constant. By using the Stokes-Einstein relation, a hydrodynamic radius is attributed to the diffusion constant. In case of the free GFP this is 5 nm and the complex has an radius of 10 nm.

Example 3: Detection of Interactions between Biomolecules and Discrimination of Nucleic Acids by Size

The thermo-optical characterization of the present invention provides the means for fast all optical biomolecule analysis. Present methods for detection and quantification of biomolecular interactions are very time consuming which means that the time needed for an analysis is on the order of 30 minutes up to hours. The present invention can detect and quantify biomolecular interactions within 1 second up to 50 seconds. The term interaction comprises interaction between biomolecules (e.g. protein, DNA, RNA, hyaluronic acids etc) but also between modified nanoparticles/micro beads and biomolecules. A typical experiment to detect/quantify interactions can be described as follows:

Step 1a, Background Measurement:

The sample buffer without fluorescently labelled sample molecules/particles is filled in the microfluidic chamber and the fluorescence is measured, while the excitation light source is turned on.

Step 1b, Determination of Fluorescence Level before Laser Heating:

An aqueous solution of a fluorescently labelled sample (e.g. biomolecules, nanoparticles, microbeads whereas all of them have a specific affinity for other biomolecules) at a given concentration is filled in a microfluidic chamber (preferably a capillary) which preferably guarantees a defined height of the chamber. Fluorescence is excited and recorded with (CCD-Camera) or without (Photomultiplier tube, Avalanche Photodiode) spatial resolution for less than 10 seconds on a CCD device or photomultiplier with exposure times of 25 milliseconds up to 0.5 seconds. Then the fluorescence excitation is turned off.

Step 2, Starting of Infrared Laser Heating:

In the following the infrared heating laser is turned on and the spatial temperature distribution is established within a few milliseconds within the solution. The temperature gradient has been calibrated once and it is not necessary to repeat this calibration every time an experiment is performed. In particular a setup were infrared heating and fluorescence imaging are performed through the same optical element from one side is advantageous for the stability of the optical and infrared foci.

It is of advantage that in the experiment the decrease of fluorescence due to photobleaching is lower than 5%.

The maximal temperature elevation is below the temperature which is known to cause damage to the molecules in the solution or disturb their interaction (e.g. temperatures between 1 and 5° C. above ambient temperature).

Depending on the thermophoretic properties of the molecules in the solution (i.e. if they move fast in a thermal gradient or slow) the infrared laser heats the solution for 5 seconds up to 100 seconds.

Step 3, Recording of the Spatial Fluorescence (i.e. Concentration) Profile:

After this period of time the fluorescence excitation is turned on and images are recorded with the same frame rate and length as described in step 1b. Step 3 is the last acquisition step necessary for evaluation of thermo-optical properties.

For detection and quantification of interactions more measurements following the protocol described previously are necessary. Step 1a is repeated with sample buffer and in step 1b the aqueous solution of a fluorescently labelled sample is mixed with an amount of the biomolecule with which the interaction should be detected or quantified. For the detection of an interaction it is necessary to mix the fluorescently labelled sample with a sufficient amount of binding partner so that a substantial amount of the fluorescently labelled molecule is in the complex with the binding partner. If the strength of the interaction should be quantified in terms of a dissociation or association constant (Ka, Kd), than the procedure described previously has to be conducted with varying concentrations of binding partner (e.g. 0.1×-10× the concentration of the fluorescently labelled binding partner). This means that a titration of binding partner should be performed.

Processing the Raw Data:

For a linear bleaching correction it is necessary to wait for the back-diffusion of all molecules following the end of step 3. This increases the time consumption of the analysis dramatically. For precise and fast measurements it is of advantage to determine the strength of bleaching from image to image and correct every individual image with its own bleaching factor. For a precise bleaching correction it is important that the temperature gradient at distance from the heat spot is low (e.g. below 0.001 K/μm). The images taken in step 1b are used to correct all images for inhomogeneous illumination. In case fluorescence is recorded without spatial resolution (e.g. avalanche photodiode or photomultiplier) photobleaching is corrected best by determining once the bleaching characteristic of a certain dye without heating laser in a control experiment.

Data Evaluation:

Qualitative detection of interaction: From the image series the spatial fluorescence distribution of the reference experiment (i.e. fluorescently labelled molecule/particle without binding partner) and the second experiment (i.e. were the binding partner is present) is extracted. The fluorescence is plotted versus the distance from the heat spot. An averaging is only possible for pixels with the same temperature and same distance. The spatial concentration distribution is obtained by correcting the fluorescence intensities for the respective temperature dependence of the dye. With knowledge of temperature dependence of the fluorescent dye and the spatial temperature distribution, the effect of a decreasing fluorescence due to temperature increase can be corrected. For the qualitative detection of interaction as well as their quantification a correction for temperature dependency is not necessary, and the spatial fluorescence distribution is sufficient. This enables us to use any fluorescent dye on the market without characterization of its temperature dependency.

The values of the fluorescence profile are integrated up to the distance were the temperature is below 10% of the maximum temperature (e.g. 70 μm). The integrated values are compared and a change give a precise indication if there is an affinity between the substances at the concentrations used, since the interaction changes the thermo-optic properties (e.g. thermophoretic mobility, surface size and chemical groups on surface). In most cases the interaction leads to higher fluorescence (concentration) at higher temperatures.

In case the whole cross-section of a capillary is heated (i.e. using cylindrical lenses to give the IR laser beam a ellipsoidal shape, which heats a cross section of a capillary homogeneously), the intensity of a lot more pixels from the centred heat spot can be averaged since all pixels at same distance to the heated line have the same temperature. This is advantageous for high precision measurements. In case fluorescence is recorded without spatial resolution the fluorescence change in the centre of the heat spot/line is measured, whereas it is again advantageous to heat the whole cross section. In general if more than a single frame is recorded in step 1b and 3 an integration of multiple frames is possible.

For a quantification of affinities the same procedure is performed for all experiments at various concentrations of non fluorescent binding partner. The result of the integration for the reference experiment (i.e. without binding partner) is subtracted from the integrated values obtained for the different concentrations of binding partners. From this evaluation on gets the amount of interacting complexes in arbitrary units. By dividing the values by the value were binding is saturated the relative amount of complexes at a certain concentration of binder is obtained. From these dataset also the concentration of free non fluorescent binding partner can be determined and the strength of the interaction can be quantified in terms of association or dissociation constant (see FIG. 25).

Example 3a: Interaction of Proteins

FIG. 25 shows a quantification of interaction between biomolecules. 100 nM of a fluorescently labelled antibody in 1×PBS buffer (anti-Interleukin 4, Sigma-Aldrich) are titrated with various amounts interleukin 4 1×PBS buffer (0-300 nM IL4 Sigma-Aldrich). (left). Approximately 200 nl of the sample mix are soaked into a capillary of 40 nm inner diameter (World Precision Instruments). The capillary is situated on a device as shown in FIG. 27. Fluid drift is prevented by closing the valves on both sides of the capillary. The measurement is performed with the device shown in FIG. 20. The spatial fluorescence distribution in steady state is measured using the previously described protocol. Three results for 5 nM, 80 nM and 300 nM are shown exemplarily. After each measurement the capillary is flushed with approximately 5 μl of 1×PBS buffer. The binding of IL 4 to the antibody changes the signal dramatically from fluorescence decrease to a fluorescence increase. Integration of the fluorescence profile up to 80 μm (distance from the heated centre, see procedure described in this example herein above) allows to determine the number of complexes in solution. (right) The concentration of free Interleukin 4 can be calculated plotted versus the concentration formed complex. These data can be fitted to determine the K_(D).

Example 3b: Discrimination of DNA by Size and Interaction of DNA Strands

Qualitative detection of interaction, using a modified protocol according to this example 3, is shown in FIG. 5. Here the concentration profiles of 20 base pair DNA hast been compared to 50 base pair DNA at a maximum temperature increase of 10° C. (at ambient temperature of 20° C.). In a second experiment 20 bases single stranded DNA is compared with 20 base pair double stranded DNA. All four experiments were performed in 1×SSC buffer the following way: 2 μl of sample are pipetted on an object slide (Roth, 1 mm thick) an sandwiched between an cover slip of 12 mm diameter. The aqueous solution spreads uniformly in between the two glass surfaces yielding a thin sheet of water with a height of approximately 20 μm. The liquid sheet is sealed by using nail polish. This avoids rapid evaporation of the sample. The microfluidic chamber is placed on the object stage of the thermo-optical setup (e.g. FIG. 1a ) and imaged via a 40× oil objective (NA 1.3, Zeiss). The laser focus is positioned so that it is approximately in the center of the field of view and has a half width of about 20 μm. Only a single pixel of a CCD camera used for detection of the fluorescence. This pixel measures the fluorescence in the center of the heat spot. The fluorescence is recorded for approx. 1 seconds without laser heating, then the IR Laser is turned on while fluorescence is still recorded. After 20 seconds of laser heating the measurement stopped. As can be seen from FIG. 5 single stranded DNA can be discriminated from double stranded DNA and DNA of different length can be discriminated within a time span of a few seconds. Bleach correction is not performed in this measurement strong change in concentration is observed.

Example 4: Detection of Binding of PEG Molecules to Nanoparticles

As mentioned previously it is also possible to detect the binding of molecules to larger inorganic particles or nanocrystals using the procedure described previously. Inorganic CdSe particles (core diameter about 12 nm) have been modified with a varying numbers (1 up to 3) of Poly-ethylen-glycol (PEG) of different molecular weight. 2 μl of sample are pipetted on an object slide (Roth, 1 mm thick) and sandwiched between a cover slip of 12 mm diameter. The aqueous solution spreads uniformly in between the two glass surfaces yielding a thin sheet of water with a height of approximately 20 μm. The liquid sheet is sealed by using nail polish. This avoids rapid evaporation of the sample. The microfluidic chamber is placed on the object stage of the thermo-optical setup (e.g. FIG. 1a ) and imaged via a 40× oil objective (NA 1.3, Zeiss). The laser focus is positioned so that it is approximately in the center of the field of view and has a half width of about 100 μm. The maximum temperature increase was determined to 5° C. above ambient temperature. The spatial fluorescence profile is recorded as described previously for the detection of biomolecular interactions. Also the raw data are processed as described previously. To measure the number or size of PEG molecules bound to the nanocrystals it is sufficient to compare the spatial fluorescence profiles obtained with the protocol described previously. However, a correction for the temperature dependent decrease of the fluorescence allows a quantification in terms of the Soret coefficient. FIG. 26 shows that the Soret coefficient increases linearly with the number of PEG molecules bound to the nanocrystals. The slope of the increase depends on the molecular weight of the PEG. FIG. 26 shows that the binding of single molecules of the size of a protein is detectable.

Example 5: Thermophoresis of Proteins

An example how the conformation, structure and surface of a molecule effect the thermo-optical characteristic of said molecules and how these characteristics may be measured, detected or characterized is given in the following:

A sample of fluorescently labelled BSA (Bovine Serum Albumin, Fermentas) is transferred in a microfluidic chamber (e.g. capillary). The temperature of the whole sample volume is adjusted by a Peltier element in thermal contact to the solution (the microfluidic chamber is placed on a device shown in FIG. 27). The Peltier element is used to regulate the “ambient temperature” of the solution. It does not create a spatial temperature distribution. The adjustment of temperature is important because the thermo-optic properties (e.g. surface properties, conformation) at varying ambient temperatures (i.e. protein conformations) should be measured. The thermo-optic properties are measured following the protocol described previously for biomolecule interactions (step 1 to step 3, without addition of binding partners since only intramolecular properties are measured). Also the processing of the raw data follows the procedure described for the detection of molecule interactions. The thermo-optical properties are evaluated by determining the concentration profile in steady state. The thermo-optic properties are plotted as Soret coefficient ST as shown in FIG. 28. The Soret Coefficient is obtained by correlating the concentration profile in an exponential fashion (c=c₀e^(−S) ^(T) ^((T-T) ⁰ ⁾) to the temperature distribution. The Soret-Coefficient is sensitive to changes in the interplay between the amino acids of the protein and the water molecules. At low temperatures the molecules are accumulated in a region of elevated temperature, which corresponds to a negative Soret coefficient. At increasing temperatures the accumulation of the molecules changes (i.e. the Soret coefficient increases). This can be readily explained by changes in the conformation of the molecule (e.g. hydrophobic groups or loops rearrange). As can be seen from FIG. 28 the sign of thermophoresis changes from negative values at lower temperatures to positive values (i.e. depletion) at higher temperatures. The sudden jump to positive Soret coefficients correlates very well with the temperature were thermal denaturation occurs (i.e. 50° C.). In the range of body temperature (i.e. 30° C.-40° C.) the thermo-optical signal does not significantly change. An explanation for this unexpected behaviour is that the protein is evolutionary designed to be functional within this temperature range. Since there is a tight structure-function relationship in nature, the structure is preserved in this temperature range. The values shown in FIG. 28 are corrected for the temperature dependence of the fluorescence dye.

The local temperature increase in the system causes a change in fluorescence which is not purely caused by changes in concentration due to thermophoresis. Since the temperature of transition from negative to positive values is important for measurements of absolute protein stability, a correction for the temperature dependence of fluorescence is advantageous. In application were differences in stability should be detected (e.g. small molecule binding to a protein) the correction for temperature dependence of the fluorescence is not necessary. The argument that protein structure/conformation is measured is supported by FIG. 28b where the experiment is started at high temperatures. Even at temperatures below the thermal denaturation temperature the Soret coefficient is positive. This is based on the slow refolding time compared to the speed of measurement, which is faster than 50 seconds. After a certain time span the thermo-optical ST values shown in FIG. 28a are also obtained again in the measurement shown in FIG. 28b . As a control the temperature of the system is increased again and the negative ST values are obtained as expected at temperatures of about 40° C.

Example 6: Detection of Conformational Changes, Like Denaturation of Proteins

-   -   a) An example where the denatured form of a protein is         distinguished from the native form without any correction for         the temperature dependence of the fluorescence dye, given in         FIG. 15. An aliquot of a fluorescently labelled Bovine Serum         Albumin is heated up to 90° C. for 10 minutes, which is well         above the temperature of denaturation and the protein cannot         refold. Also the thermo-optic characteristics of a native         aliquot (i.e. not heated to 90° C.) is measured. Starting with         the native sample, approximately 200 nl of the sample are soaked         into a capillary of 40 nm inner diameter (World Precision         Instruments). The capillary is situated on a device as shown in         FIG. 27. Fluid drift is prevented by closing the valves on both         sides of the capillary. The measurement is performed with the         device shown in FIG. 20. The spatial fluorescence distribution         in steady state is measured using the previously described         protocol. After each sample the capillary is flushed with         approximately 5 μl of 1×PBS buffer. The experiment is performed         as described for the measurement of biomolecular interactions         (Step 1-Step 4). The experiment is performed twice with         different infrared laser powers (i.e. maximal temperatures of         5° C. and 10° C. above ambient temperature are employed).         Afterwards the sample with the denatured protein is measured.         Again three experiments with different laser powers are used. In         FIG. 29 the fluorescence is plotted as a function of the         distance to the heat source (i.e. laser focus) is shown for the         native and denatured form at two different laser powers (i.e.         maximum temperatures of 5° C. and 10° C.). In both cases there         are two contributions to the fluorescence change. First there is         an increase or decrease in concentration (for the native or         denatured form, respectively) and secondly there is a decrease         in fluorescence due to the temperature dependence of the         fluorescence dye. For a qualitative comparison (i.e. to         distinguish between a native and denatured form.) a correction         for the temperature dependence of the dye is not necessary. In         all cases shown in FIG. 29 the fluorescence decreases. But as         expected the decrease in fluorescence for the native protein is         not as strong decrease observed for the denatured form. This is         readily explained by the negative Soret coefficient of the         native protein at 20° C. ambient temperature (see also FIG. 28),         which leads to an accumulation of molecules at higher         temperatures. This counteracts the decrease in fluorescence         caused by the temperature dependence of the fluorescence. At         higher laser powers (i.e. a maximum temperature increase of 10°         C.) the difference between native and denatured form is even         stronger because the thermophoretic accumulation and         thermophoretic depletion, for the respective form, get stronger.         Also smaller conformational changes can be detected on the basis         of different strength of thermophoresis. A different direction         of thermophoretic movement is advantageous but not necessary.     -   b) A sample of fluorescently labelled bovine serum albumin (BSA)         has been split in two parts. One is only exposed to ambient         temperatures, while the other half is heated up to 100° C. for         several minutes (i.e. irreversibly denatured). The         thermo-optical properties of both samples (native and denatured)         are measured at 800 mA power of the infrared laser (i.e. maximal         temperature increase of 20° C.). As can be seen from FIG. 30,         the fluorescence of the denatured protein is lower than the         fluorescence of the native protein. This is explained as         follows. The fluorescence dye of both samples shows the same         decrease in fluorescence due to the increase in temperature         (i.e. temperature sensitivity of the fluorescence). But the         denatured protein shows a positive thermophoretic mobility (i.e.         moves to the cold), while the native protein has a negative         thermophoretic mobility (i.e. moves to the hot). The         accumulation at elevated temperatures is the reason, why the         decrease in fluorescence is lower for the native protein, while         the denatured protein is, in addition to temperature dependency,         depleted from the region of elevated temperature. Interestingly,         by approaching the denaturing temperature (i.e. 50° C.) of the         protein the amplitudes of the native and denatured protein         approach each other an are essentially the same. This means that         by measuring the amplitude of the fluorescence change an         comparison to the reference sample allows to detect the melting         temperature of a protein and to discriminate between native and         denatured form of a protein. And to detect a shift in melting         temperature due to interactions of the protein with other         biomolecules or small molecules (e.g. drug candidates).

Example 7: Optothermal Trap/Thermooptical Trap

In the following, silica particles are employed as illustrative particles/beads to be thermo-optically trapped by the method of the present invention. It is understood that the described method can also be employed for the thermo-optical trapping of other molecules, like biomolecules or lipid vesicles (as also illustrated in a further example).

Silica particles (1 μm diameter, plain, Kisker Biotech) are diluted 1/100 in distilled water. 2 μl are pipetted on an object slide (Roth, 1 mm thick) an sandwiched between an cover slip of 12 mm diameter. In the following the term bead is used as a synonyme for particle. The aqueous bead-containing solution spreads uniformly in between the two glass surfaces, yielding a thin sheet of water with a height of approximately 20 μm. The liquid sheet is sealed by using nail polish. This avoids rapid evaporation of the sample. The microfluidic chamber is placed on the object stage of the thermo-optical setup (as e.g. illustrated in the appended FIG. 1a ) and imaged via a 40× oil objective (NA 1.3, Zeiss). The laser focus is positioned so that it is approximately in the center of the field of view and has a half width of about 100 μm. Then the IR Laser is turned on and heats the solution to 10° C. above ambient temperature (20° C.) at the maximum of the spatial temperature distribution in the center of the IR laser focus. The images series shown in FIG. 33 illustrates the process of particle accumulation/trapping. In the beginning (first image of FIG. 33, top of the page), without laser heating, the beads are almost equally distributed. The black circle shows the position of the laser focus. The following images show the development of the particle distribution in the next three seconds after the heating laser is turned on. The particles show a directed movement to the region of elevated temperature at the laser focus, which can be also referred to as negative Soret effect or negative thermophoresis. Surprisingly, silica particle show negative thermophoresis at room temperature. The particles are trapped in the center of the temperature distribution at the laser focus. The particle experiences a potential well created by the spatial temperature distribution. The directed motion to the region of highest temperature is explained by the particle's tendency to minimize its energy of solvation. The position of the particle is not exactly in the center of the heat spot, since the thermal fluctuations push the particle out of its position. Appended FIG. 34 illustrates that the beads are trapped within the region of highest temperature even when the stage (i.e. sample) is moved with speeds of millimeters/second relative to the fixed laser focus. By using the thermo-optical trap, particle can be moved arbitrarily and can also be concentrated. After confining antibody modified beads to the heat spot, an interaction between the particles due to a binding of a single antigen in the solution to more than one bead can be detected.

Another approach of a thermo-optical characterization in accordance with this invention is shown in FIG. 32. Silica particles (1 μm diameter, plain, Kisker Biotech) are diluted 1/1000 in distilled water. The dilution factor is empirical. The dilution is such that only a single particle is observed in a region of approximately 400 μm times 400 μm. 2 μl are pipetted on an object slide (Roth, 1 mm thick) an sandwiched between an cover slip of 12 mm diameter. The aqueous bead-containing solution spreads uniformly in between the two glass surfaces yielding a thin sheet of water with a height of approximately 20 μm. The liquid sheet is sealed by using nail polish. This avoids rapid evaporation of the sample. The microfluidic chamber is placed on the object stage of the thermo-optical setup (e.g. FIG. 1a ). The laser focus is positioned so that it is approximately in the center of the field of view and has a half width of about 100 μm. Then the IR Laser is turned on and heats the solution to 10° C. above ambient temperature (20° C.) at the maximum in the center of the IR laser focus. Due to the high dilution, a single particle is trapped in a potential well created by a spatial temperature distribution (s. FIG. 32a ). As silica particles show a negative thermophoresis the well is deepest at high temperatures (i.e the particles minimize their energy of solvation at high temperatures). The single silica particle fluctuates in the potential well since the thermal fluctuations push the particle out of its position. The fluctuations are recorded via a CCD camera (at t=1 s, 2 s, 3 s, 4 s, 5 s, 6 s, 7 s) and the positions are tracked by Software (selfmade Software, Labview National Instruments, detecting the pixel with highest intensity) with nanometer resolution (see FIG. 32b ). A histogram is calculated from the positional information (see FIG. 32c ). The width of the distribution is very sensitive to the thermo-optical properties of the particle. If molecules bind to the surface of the particle, the effective potential for the bead changes and the amplitude of the fluctuations increases or decreases. By observing the amplitude change over time, a kinetic binding curve can be measured. In a microfluidic system such an experiment is performed as follows: a solution containing a 1/1000 dilution (distilled water) is flushed or soaked into a capillary (see FIG. 27). The valves at the end of the capillary are closed. A single particle (modified (e.g. coated) with an antibody specific to a certain antigen, e.g. Interleukin 4) is trapped and the fluctuations of the particle are detected for 10 seconds to 100 seconds. The beads is surrounded by pure buffer solution. In a next step the buffer is exchanged with a buffer containing the respective antigen. While the buffer is exchanged the bead is still trapped. After exchange of buffer the fluctuation of the same bead as before are recorded. The change in fluctuation amplitude is used to detect interactions between antibody and antigen. Its change over time is used to measure binding kinetics.

By using a device as illustrated in the appended FIG. 24, a temperature gradient can be generated in solution, by scanning lines (e.g. 10) perpendicular to each other in the solution. Where these lines cross, a temperature maximum is observed. Points on the scanned line have an intermediate temperature, while spaces in between the lines represent temperature minima (e.g. ambient temperature if spaces in between the lines are sufficiently wide). The silica particle described above will move to equilibrium positions on the crossing points of the heated lines. By moving the temperature grating all particles will move simultaneously. In addition the fluctuation of all beads may be measured simultaneously.

Example 8: DNA Melting Curves

Standard protocol for the measurement of melting curves (using the device shown in e.g. FIG. 1a, 1b , 16 to 18, 20 to 24 or 37):

-   -   Raman-Laser 1455 nm; coupled to galvanometric mirrors via fibre     -   No collimator after fibre, Laser-beam hits mirrors divergently;         mirrors are reflecting the laser-beam onto a lens; laser beam is         focused to the chamber via the lens.     -   Laser on/off is controlled via moving the laser in an out of the         field of view via the mirrors.

Preparations:

-   -   Coverslips (170 μm) thick, one with a diameter of 12 mm the         other quadratic 24×24 mm rinsed with deionized water, rinsed         with ethanol, then again rinsed with deionized water.     -   Dilution of solutions:         -   10 μM Tamra in 1×SSC         -   Hairpin from 100 μM in MilliQ-water diluted to 10 μM, 1 μM,             100 nM, or less in SSC-Buffer (1×, 0.5×, 0.1× or less).         -   add the detergent TWEEN20 to an end volume of 0.01% (only if             there is unspecific adsorption).

Adjustment:

-   -   Check if everything is ok with the microscope (apertures,         filters)     -   10 μM Tamra (tetramethylrhodamine) in 1×SSC (150 mM NaCl, 15 mM         Na3-citrate, pH 8.1), with 0.01% TWEEN 20; volume of 2 μl into         chamber built with 2 Coverslips (170 μm thick), sealed with nail         polish     -   Wait until nail polish is dry     -   Add immersion oil onto the upper coverslip     -   Put chamber onto the measurement stage     -   Focussing fluorescence image by nearly closing the aperture     -   Find laser spot     -   Focus laser     -   Defocus laser by increasing the distance between lens and         chamber     -   Adjust laser focus in a way that a broad and appropriate         temperature distribution is established     -   Move laser spot with galvanometric mirrors out of the field of         view in such a way that the influence of the laser to the         fluorescence image is minimal→save this settings to the         measurement program (avoid to cross the zero point on the         voltage scale of the galvanometric mirrors)     -   Measure the room temperature

Measurement:

-   -   Use the following steps for the measurement of temperature as         well as for the measurement of melting curves:     -   Conduct the measurement with the trigger-programme.     -   Settings: 40× oil immersion-objective, 8×8 Binning, 10 ms         exposure time (→28 Hz readout rate)     -   To do manually:     -   1. Laser on     -   2. Light source on (open shutter in case of HXP), write down the         settings of the light source     -   3. Start the trigger program     -   4. Laser off     -   Light source on

The measurement is based on the detection of fluorescence and therefore may be conducted in a device according to appended FIGS. 16-18 and/or 20-24. As the exact timing of measurements is preferable, the used devices like CCD, IR-Laser and light source are synchronised by the use of an electronic trigger signal. As the used CCD-camera (Andor Luca) has a trigger output port, the IR-Laser control element and the light source control element are synchronised with the trigger signal of the CDD-camera. In particular the second high level of the CCD trigger output signal is taken as the zero point of time of the measurement. In the following the exposure time of the CCD-camera is 10 ms and the minimal time span between two images is determined by the frame rate of the CCD-camera. As chamber a 2 μl solution of the fluorescently labelled sample molecules (e.g. 10 μM tetramethylrhodamine (TAMRA) in 75 mM NaCl, 7.5 mM Na3-citrate, 0.01% TWEEN 20, pH 8.1) is sandwiched between two 170 μm thick glass coverslips with diameter of 12 mm and sealed with nail polish. This results in a thickness of the chamber of about 20 μm. Then the chamber is moved to the measurement device and the optics are focussed to the chamber.

Brief Description of the Sequence of the Measurement:

Before the measurement starts, the fluorescence background of the measurement device is recorded in “Step 0”. As this background is characteristic for the used device and may not change during a long time period this step is preferably done only once to characterize the used device.

Time t = 0: Step 1: a first fluorescence image of the spatial fluorescence distribution in the chamber is recorded. Time t = 20 ms Step 2: the IR-Laser is switched on Time t = 60 ms Step 3: a second fluorescence image of the spatial fluorescence distribution in the chamber is recorded.

After these steps the measurement is finished and the raw data processing and the data evaluation is conducted.

Detailed Description of the Sequence of the Measurement:

Step 0, Background Measurement:

A sample buffer without fluorescently labelled sample molecules/particles is filled into a microfluidic chamber and the spatial fluorescence distribution in the chamber is measured with the CCD-camera, while the excitation light source is turned on.

Step 1, Determination of Fluorescence Level at Ambient Temperature before Laser Heating:

An aqueous solution of a temperature sensitive dye (e.g. TAMRA) at a given concentration (e.g. 10 μM) is filled in the chamber which preferably provides a defined height of the chamber. Fluorescence is excited with the light source (LED) and recorded with spatial resolution with the CCD-camera at ambient temperature.

First the CCD-camera is started and the first high level of the camera trigger output signal is used to synchronise the IR-Laser control element and LED control element with the CCD-camera. For the synchronising a measurement card from e.g. National Instruments can be used.

For good illumination the LED is turned on with the use of the camera trigger signal a short time before the CCD-camera records a first fluorescence image I₀(x,y). Therefore during the exposure time of the CCD-camera of 10 ms the fluorescence excitation light source has reached its steady state level of light output. When the camera starts its recording the output trigger signal of the camera is the second time at the high level state. This second high level state of the output signal determines the zero point of time t=0.

Step 2, Starting of Infrared Laser Heating:

The infrared heating laser is turned on at time t=20 ms and the spatial temperature distribution is established within a few milliseconds within the solution. The temperature distribution has been calibrated once in a way that for example all temperatures between 30° C. and 90° C. are present in the recorded image and it is not necessary to repeat this calibration every time an experiment is performed.

Step 3, Recording of the Spatial Fluorescence Profile with Infrared Laser Heating:

At t=60 ms, 40 ms after IR-Laser irridation was started, a second fluorescence image I₁(x,y) is recorded with the CCD-camera with an exposure time of 10 ms.

After the CCD-camera has taken the second picture, the first and the second picture are saved to the hard disk of a PC for processing the raw data and for data evaluation. Then the measurement process is finished.

Processing the Raw Data and Data Evaluation:

Because of the short time of measurement no correction for bleaching may be necessary. The images I₀ and I₁ are corrected against camera background. Then calculating the ratio K(x,y)=I₁(x,y)/I₀(x,y) for each pixel of the fluorescence images (second image divided by the first image, both background corrected) ensures the removal of artefacts from an inhomogeneous illumination. As the temperature dependence F(T) (FIG. 15) of the dye TAMRA is known from a calibration experiment in a fluorimeter, the spatial temperature distribution T(x,y) (FIG. 3c ) can be derived from to the ratio K(x,y).

As the timing of measurement is the same for the temperature measurement and for the measurement of the melting curve, both measurements can be conducted at one time in one chamber if the emission spectrum of the dye for the temperature measurement (e.g. TAMRA) can be well separated from the emission spectrum of the fluorescent label of for example the molecular beacon (e.g. HEX as fluorophore (see appended FIG. 6) and Dabcyl as a quencher).

Example 9: Detection of Covalent and Non-Covalent Modifications of Nanoparticles

An example for the detection of covalent and non-covalent modification is shown in FIG. 35. Nanoparticles (i.e. nanocrystals or quantum dots) have been obtained from Invitrogen. Particles were purchased which have charged polymere coating to be stable in aqueous solution (diameter 12 nm). In addition particles with covalently coupled strepavidin were purchased (diameter 21 nm), as well as a biotinylated 40 bases single stranded DNA. The following samples have been prepared: First unmodified nanoparticles diluted to 1 μM concentration in 1×SSC (Saline-Sodium Citrate) buffer. Secondly, streptavidin coated nanoparticles diluted in 1×SSC buffer to 1 μM concentration. Thirdly, streptavidin coated nanoparticles diluted in 1×SSC buffer to 2 μM concentration were mixed with 2 μM 40 bases biotinylated single stranded DNA (IBA GmbH, Göottingen) in 1×SSC buffer in a 1/1 ratio. All three experiments were prepared in 1×SSC buffer in accordance with the following protocol: 2 μl of sample are pipetted on an object slide (Roth, 1 mm thick) an sandwiched between an cover slip of 12 mm diameter. The aqueous solution spreads uniformly in between the two glass surfaces leading to a thin sheet of water with a height of approximately 20 μm. The liquid sheet is sealed with nail polish to avoid rapid evaporation of the sample. The microfluidic chamber is placed on the object stage of the thermo-optical setup (e.g. FIG. 1a ) and imaged via a 40× oil objective (NA 1.3, Zeiss). The laser focus is positioned so that it is approximately in the center of the field of view and has a half width of about 20 μm. The experiments were conducted as described herein above for the detection of interactions, with a maximum temperature increase of 5° C. above room temperature. A correction for temperature dependent fluorescence has been performed for precise measurement of the Soret Coefficient from the spatial concentration distribution. (i.e. the temperature dependency of the fluorescence of the nanocrystal has been determined in an independent fluorimeter experiment). As can be seen from FIG. 35, the Soret coefficient is positive and larger for the streptavidin coated nanoparticle than for the nanoparticle without protein. Furthermore the binding of a single single-stranded DNA molecule to the particle is detected as a change in Soret coefficient. This is interesting, since the short flexible DNA molecule does not contribute substantially to the size of the nanoparticle (as the streptavidin does). Since thermophoresis is sensible to changes of the surface properties upon binding of the DNA molecule, the binding of the comparatively small DNA molecule to the particle can be detected.

Example 10: Thermophoresis and Thermophoretic Trapping of Lipid Vesicles

Texas-Red DHPE (1,2-dihexadecanoyl-sn-glycero-3-phosphoethanolamine, triethylammonium salt), a fluorescently labelled phospholipid, is used for staining the vesicles. The lipid is added to the process of vesicle formation with approximately 1 mol percent with respect to the main constituting phospholipids. The following stock solutions were used for preparation of the vesicles by electro swelling:

Lipid Stock Solution:

-   -   2.5 mg/ml DPhPC (1,2 Diphytanoyl-sn-Glycero-3-Phosphocholine) in         Chloroform (CHCl3)     -   1-2% stearylamine         -   1% fluorescent lipid             (1,2-dihexadecanoyl-sn-glycero-3-phosphoethanolamine,             triethylammonium salt (Mol. Probes)

Sucrose Stock Solution:

-   -   300 mM Sucrose in distilled water

Buffer Solution:

-   -   150 mM KCl, 20 mM MES, pH 5

The final vesicle solution is diluted 1/100 in distilled water. 5 μl of the dilution were pipetted on an object slide. The droplet was sandwiched in between the object slide and a round 12 mm diameter coverslip and placed on the measurement apparatus. Fluorescence was observed through an oil objective using the setup from 1 a, and alternatively the setup from FIG. 19. In the latter case, IR-Laser heating is performed through the same objective. The maximum temperature increase was about 15° C. The temperature profile had a half width of 20 μm. Images before and after infrared laser heating were taken. Two representative images before and 10 seconds after infrared laser heating are shown in FIG. 39. The vesicles are attracted by heat spot due to negative thermophoresis. They are accumulating in the centre of the heat spot. The catchment area is dependent on the width of the temperature profile (i.e. a sufficient strong temperature gradient is necessary to drive directed particle motion toward the heated centre within a finite time). The trapping of vesicles has several, important applications, e.g. vesicles (as well as cells) can be transported and moved within a solution. Also, the fluctuation of single vesicles in the potential well (i.e. created by temperature increase) can be observed. Since the amplitude of fluctuations is dependent on the properties of the vesicle, any changes, like protein binding to vesicles or the activity of a membrane protein, e.g. an ion-pumping membrane protein can be detected as a change in fluctuation amplitude. The sign of the thermophoretic motion (e.g. attracted to the heat centre or repelled by the heated centre) is dependent on the properties (e.g. charge, size, surface modification, protein binding). When a vesicle which is normally attracted by the region of highest temperature is repelled from the heated centre, this is indicative for a change in the properties of this vesicle. This behaviour may be observed after changing the buffer around the vesicle to a solution containing, e.g. a binding partner. Also the behaviour of two samples containing vesicles in buffer and vesicles in buffer with binding partner (or membrane protein activating substance, e.g. ATP) may be compared or observed. Finally, the differences in the thermophoretic properties can be used to sort vesicles or cells. 

The invention claimed is:
 1. A method for measuring the stability of a biomolecule in a solution comprising: (a) providing a sample with the biomolecule in a solution in a capillary or multi-well plate; (b) exciting fluorescently the biomolecule and detecting a first fluorescence of the excited biomolecule; (c) irradiating a laser light beam into the solution to obtain a spatial temperature distribution in the solution around the irradiated laser light beam; (d) detecting a second fluorescence of the biomolecule in the solution at a predetermined time within the range of 1 ms to 250 ms after irradiation of the laser into the solution has been started, and (e) characterizing the stability of the biomolecule based on the first and second detections.
 2. The method of claim 1, wherein unfolding of the biomolecule is detected and the stability of the biomolecule is characterized based on the detected unfolding.
 3. The method of claim 1, wherein the biomolecule is a protein or a peptide.
 4. The method of claim 3, wherein the protein is an antibody or a fragment or derivative thereof.
 5. The method of claim 3, wherein the fluorescence of tryptophan, tyrosine or phenylalanine in the biomolecule is detected.
 6. The method of claim 1, wherein the biomolecule is a nucleic acid.
 7. The method of claim 1, wherein both said fluorescence detection and said laser irradiating are performed through a single optical unit.
 8. The method of claim 7, wherein both fluorescence detection and laser irradiating are performed through a single objective.
 9. The method of claim 1, wherein the detection time is in the range of 1 ms to 50 ms.
 10. The method according to claim 1, wherein the spatial temperature distribution in the solution around the irradiated laser light beam is 0.0 to 5 K/μm.
 11. The method of claim 10, wherein the spatial temperature distribution in the solution around the irradiated laser is 0.0 to 2 K/μm.
 12. The method of claim 1, wherein the laser beam is irradiated through an optical element into the solution.
 13. The method of claim 12, wherein the optical element is a single lens.
 14. The method of claim 1, wherein said fluorescence is detected with a CCD camera, a photodiode or a photomultiplier tube.
 15. The method of claim 1, wherein the laser light has a wavelength of 1200 nm to 2000 nm.
 16. The method of claim 1, wherein the spatial temperature distribution is between 0.1° C. and 100° C.
 17. The method of claim 1, wherein the temperature gradient is created within 0.1 μm to 500 μm in diameter around the laser beam.
 18. The method of claim 1, wherein the detection of the fluorescence is detected within a range of from 1 nm to 500 μm in the direction of the laser beam. 